Extended or continuous wear silicone hydrogel contact lenses for the extended release of comfort molecules

ABSTRACT

A drug delivery system is disclosed. The drug delivery system includes a recognitive polymeric hydrogel through which a drug is delivered by contacting biological tissue. The recognitive polymeric hydrogel is formed by using a bio-template, which is a drug or is structurally similar to the drug, functionalized monomers, preferably having complexing sites, and cross-linking monomers, which are copolymerized using a suitable initiator. The complexing sites of the recognitive polymeric hydrogel that is formed preferably mimics receptor sites of a target biological tissue, biological recognition, or biological mechanism of action. The system in accordance with an embodiment of the invention is a contact lens for delivering a drug through contact with an eye.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application claims priority under 35 U.S.C. §119 to provisional application Ser. No. 61/598,036 filed Feb. 13, 2012, titled “Extended or Continuous Wear Silicone Hydrogel Contact Lenses For the Extended Release of Comfort Molecules,” herein incorporated by reference in its entirety.

This application is also a Continuation-in-Part of U.S. Ser. No. 12/395,394 filed Feb. 27, 2009, titled “Controlled and Extended Delivery of Hyaluronic Acid and Comfort Molecules Via a Contact Lens Platform,” which claims the priority under 35 U.S.C. §119 of the U.S. Provisional Application Ser. No. 61/067,261 filed Feb. 27, 2008, titled “Controlled and Extended Delivery of Hyaluronic Acid Via a Contact Lens Platform,” The application Ser. No. 12/395,394 is also a Continuation-in-Part of U.S. Ser. No. 11/985,262, filed Nov. 13, 2007, now U.S. Pat. No. 8,349,352 issued on Jan. 8, 2013, titled “Therapeutic Contact Lenses with Anti-Fungal Delivery,” which claims priority under 35 U.S.C. §119 of U.S. Provisional Application Ser. No. 60/858,584 filed Nov. 13, 2006. The application Ser. No. 12/395,394 is also a Continuation-in-Part of U.S. Ser. No. 11/346,770, filed Feb. 3, 2006, now U.S. Pat. No. 8,349,351 issued on Jan. 8, 2013, which claims the priority under 35 U.S.C. §119 of the U.S. Provisional Application Ser. No. 60/692,049, titled “Sustained Ophthalmic Drug Delivery Via Biomimetic Recognitive Contact Lens”, filed Jun. 17, 2005; the U.S. Provisional Application Ser. No. 60/736,140, titled “Sustained Ophthalmic Drug Delivery Via Biomimetic Recognitive Contact Lens”, filed Nov. 10, 2005; and the U.S. Provisional Application Ser. No. 60/650,450, titled “Enhanced Loading and Extended Release Contact Lens for Histamine Antagonist Drug Ketotifen”, filed Feb. 4, 2005, all of which are hereby incorporated by reference.

This application is also Continuation-in-Part of U.S. Ser. No. 13/328,836 filed Dec. 16, 2011, which is a Divisional application of U.S. Ser. No. 11/346,770 filed Feb. 3, 2006, now U.S. Pat. No. 8,349,351 issued on Jan. 8, 2013, which claims the priority under 35 U.S.C. §119 of the U.S. Provisional Application Ser. No. 60/692,049, titled “Sustained Ophthalmic Drug Delivery Via Biomimetic Recognitive Contact Lens”, filed Jun. 17, 2005; the U.S. Provisional Application Ser. No. 60/736,140, titled “Sustained Ophthalmic Drug Delivery Via Biomimetic Recognitive Contact Lens”, filed Nov. 10, 2005; and the U.S. Provisional Application Ser. No. 60/650,450, titled “Enhanced Loading and Extended Release Contact Lens for Histamine Antagonist Drug Ketotifen”, filed Feb. 4, 2005, all of which are hereby incorporated by reference.

FIELD OF THE INVENTION

The invention relates to drug delivery systems. More specifically, this invention relates to systems for and method of time released ophthalmic drug delivery using contact lenses.

BACKGROUND OF THE INVENTION

Delivering medications via contact lenses has been a prevailing notion since the inception of using hydrophilic, crosslinked polymer gels on the surface of the eye. In fact, the first patent in the field from Otto Wichterle in 1965 states that “bacteriostatic, bacteriocidal or otherwise medicinally active substances such as antibiotics may be dissolved in the aqueous constituent of the hydrogels to provide medication over an extended period, via diffusion.” However, there is evidence that this notion of a dissolved component in an aqueous constituent has been around for a much longer period of time. Evidence exists that honey soaked linen was used in ancient Rome as an ophthalmic dressing in the treatment of disease.

The biggest obstacle to using the fluid entrained in the aqueous portion of the polymer gel is maintaining a significant concentration of drug within the fluid to have a therapeutically relevant effect, which is ultimately limited by the solubility of the drug. This has been the primary reason why drug release from contact lenses has not become a clinical or commercial success. To an equivalent extent, the control over the drug delivery profile and an extended release profile is also important to therapeutic success and has not been demonstrated using these methods. Drug uptake and release by conventional (i.e., currently available) soft contact lenses can lead to a moderate intraocular concentration of drug for a very short period of time, but does not work very well due to a lack of sufficient drug loading and poor control of release. The use of soft, biomimetic contact lens carriers (i.e., recognitive polymeric hydrogels) described herein has the potential to greatly enhance ocular drug delivery by providing a significant designed and tailorable increase in drug loading within the carrier as well as prolonged and sustained release with increased bioavailability, less irritation to ocular tissue, as well as reduced ocular and systemic side effects.

The ocular bioavailability of drugs applied to the eye is very poor (i.e., typically less than 1-7% of the applied drug results in absorption with the rest entering the systemic circulation). Factors such as ocular protective mechanisms, nasolacrimal drainage, spillage from the eye, lacrimation and tear turnover, metabolic degradation, and non-productive adsorption/absorption, etc., lead to poor drug absorption in the eye. Currently, more efficient ocular delivery rests on enhancing drug bioavailability by extending delivery and/or by increasing drug transport through ocular barriers (e.g., the cornea—a transparent, dome-shaped window covering the front of the eye; the sclera—the tough, opaque, white of the eye; and the conjunctiva—a mucous membrane of the eye with a highly vascularized stroma that covers the visible part of the sclera). A topically applied drug to the eye is dispersed in the tear film and can be removed by several mechanisms such as:

-   -   (i) irritation caused by the topical application, delivery         vehicle, or drug which induces lacrimation leading to dilution         of drug, drainage, and drug loss via the nasolacrimal system         into the nasopharynx and systemic circulation (e.g., the rate         drainage increases with volume);     -   (ii) normal lacrimation and lacrimal tear turnover (16% of tear         volume per minute in humans under normal conditions);     -   (iii) metabolic degradation of the drug in the tear film;     -   (iv) corneal absorption of the drug and transport;     -   (v) conjunctival absorption of the drug and scleral transport;     -   (vi) conjunctival ‘non-productive’ absorption via the highly         vascularized stroma leading to the systemic circulation; and     -   (vii) eyelid vessel absorption leading to systemic circulation.         Therefore, due to these mechanisms, a relatively low proportion         of the drug reaches anterior chamber ocular tissue via         productive routes such as mechanisms (iv) and (v).

For posterior eye tissue and back of the eye diseases (e.g., age-related macular degeneration, retinal degeneration, diabetic retinopathy, glaucoma, retinitis pigmentosa, etc.), the amount of drug delivered can be much less compared to front of the eye disease. To treat back of the eye disease, four approaches have typically been used, topical, oral (systemic delivery), intraocular, and periocular delivery.

Topically applied drugs diffuse through the tear film, cornea/sclera, iris, ciliary body, and vitreous before reaching posterior tissues, but due to the added transport resistances do not typically lead to therapeutically relevant drug concentrations. However, researchers have shown that topically applied drugs do permeate through the sclera by blocking corneal absorption and transport. Intravitreal injections (injections into the eye) require repeated injections and have potential side effects (hemorrhage, retinal detachment, cataract, etc.) along with low patient compliance. Extended release devices have been used but require intraocular surgery and often have the same incidence of side effects. Periocular drug delivery is less invasive and also requires injections or implant placement for predominantly transscleral delivery.

To overcome most of these protective mechanisms, topical formulations have remained effective by the administration of very high concentrations of drug multiple times on a daily basis. For a number of drugs high concentrations can lead to negative effects such as burning, itching sensations, gritty feelings, etc., upon exposure of the medication to the surface of the eye as well as increased toxicity and increased ocular and systemic side effects. However, traditional ophthalmic dosage forms such as solutions, suspensions, and ointments account for 90% of commercially available formulations on the market today. Solutions and suspensions (for less water soluble drugs) are most commonly used due to the ease of production and the ability to filter and easily sterilize. Ointments are used to much lesser extent due to vision blurring, difficulty in applying to the ocular surface, and greasiness. The term “eye drops” herein is meant to refer to all topological medications administered to a surface of the eye including but not limited to solutions, suspensions, ointments and combination thereof. In addition to the aforementioned problems, drug delivery through the use of eye drops does not provide for controlled time release of the drug. Eye drops medications typically have a low residence time of the drug on the surface of the eye.

The efficacy of topical solutions has been improved by viscosity enhancers that increase the residence time of drugs on the surface of the eye, which ultimately lead to increased bioavailability as well as more comfortable formulations. Also, inclusion complexes have been used for poorly soluble drugs, which increase solubility without affecting permeation.

Other recent delivery methods have included in situ gel-forming systems, corneal penetration or permeation enhancers, conjunctival muco-adhesive polymers, liposomes, and ocular inserts.

Ocular inserts, in some cases, achieve a relatively stable or constant, extended release of drug. For example, ocular inserts such as Ocusert® (Alza Corp., FDA approved in 1974) consist of a small wafer of drug reservoir enclosed by two ethylene-vinyl acetate copolymer membranes, which is placed in the corner of the eye and provides extended release of a therapeutic agent for approximately 7 days (i.e., pilocarpine HCL, for glaucoma treatment reducing intraocular pressure of the eye by increasing fluid drainage). Lacrisert (Merck) is a cellulose based polymer insert used to treat dry eyes. However, inserts have not found widespread use due to occasional noticed or unnoticed expulsion from the eye, membrane rupture (with a burst of drug being released), increased price over conventional treatments, etc.

Mucoadhesive systems and in-situ forming polymers typically have problems involving the anchorage of the carrier as well as ocular irritation resulting in blinking and tear production. Penentration enhancers may cause transient irritation, alter normal protection mechanisms of the eye, and some agents can cause irreversible damage to the cornea.

The novel soft, biomimetic contact lens carriers proposed in this work will provide a significant increase in drug loading within the gel as well as prolonged and sustained release. This will lead to prolonged drug activity and increased bioavailability, reduced systemic absorption, reduced ocular and systemic side effects, and increased patient compliance due to reduced frequency of medication and reduced irregularity of administration (i.e., eye drop volume depends on angle, squeeze force, etc., and has been experimentally verified to be highly variable). They will also be able to be positioned easily as well as easily removed with or without use to correct vision impairment. Since they will be positioned on the cornea, this will lead to enhanced corneal permeability as well.

SUMMARY OF THE INVENTION

The present invention is directed to a drug delivery methods and systems. The drug delivery system includes a recognitive polymeric hydrogel through which a drug is delivered by contacting biological tissue. The recognitive polymeric hydrogel is formed using a bio-template, which is a drug or is structurally similar to the drug, functionalized monomers, preferably having complexing sites, and cross-linking monomers, which are copolymerized using a suitable initiator, such as described in detail below. The complexing sites of the recognitive polymeric hydrogel that is formed preferably mimics receptor sites of a target biological tissue, biological recognition, or biological mechanism of action. The system unitizes what is referred to herein as a biomimetic recognitive polymeric hydrogel.

A system in accordance with an embodiment, is an ophthalmic drug system. The ophthalmic drug system includes soft contact lenses formed from the biomimetic recognitive polymeric hydrogel and that are impregnated with a drug that can be release over a duration of time while in contact with eyes. The invention is directed to both corrective or refractive contact lenses and non-corrective or non-refractive contact lenses. While the invention as described herein refers primarily to ophthalmic drug systems, it is understood that the present invention has applications in a number of different contact drug delivery systems. For example, the biomimetic recognitive polymeric hydrogel can be used in bandages, dressings, and patch-type drug delivery systems to name a few.

In an aspect of the invention, a comfort agent delivery system includes: a contact lens, the contact lens comprising a recognitive weakly cross-linked polymeric hydrogel matrix with complexing memory sites that complex a comfort agent and release the comfort agent from the hydrogel matrix over time while in contact with a surface of an eye. In an aspect, the weakly cross-linked polymeric hydrogel is formed by the steps of generating a solution comprising amounts of a bio-template, a functionalized monomer or macromer, complexing the functionalized monomer or macromer and the bio-template through non-covalent interactions, initiating copolymerization of the functionalized monomer or macromer and the cross-linking monomer, and loading the comfort agent into the memory site. In a further aspect, the loading of the comfort agent (or other drug) onto the delivery system may occur using a wash step, such as a portion of the bio-template can be washed from the recognitive hydrogel polymer, and loaded/reloaded with a comfort agent (or other drug). In a still further aspect, where the bio-template is the comfort agent (or other drug), the washing step can be illuminated. In a further aspect, the weakly cross-linked polymeric hydrogel provides controlled release of the comfort agent and optionally a drug for up to at least 30 days.

In an aspect of the invention, a drug delivery system includes: a contact lens, the contact lens comprising a weakly cross-linked hydrogel matrix, wherein the matrix has complexing sites that complex a drug and release the drug from the weakly cross-linked hydrogel matrix over time while in contact with a surface of an eye, wherein the weakly cross-linked hydrogel matrix is formed by generating a solution comprising amounts of a bio-template, a functionalized monomer and a cross-linking monomer and initiating copolymerization of the functionalized monomer and the cross-linking monomer, and loading the hydrogel matrix with the drug. In other aspects, the steps of washing the bio-template from the polymer and loading/reloading of the drug can take place. In a further aspect, the functionalized monomer and the bio-template are complexed in the solution through non-covalent interactions. In a still further aspect, the functionalized monomer contains functional groups that mimic receptor sites of a target biological tissue that are associated with a biological mechanism of the drug at that biological tissue.

In a further aspect of the invention, a method for making a drug delivery system includes: identifying receptor sites at a target biological tissue that are associated with a biological mechanism of a drug at the target biological tissue; synthesizing or selecting a functionalized monomer with functional groups that mimic the receptor sites of the target biological tissue to form a matrix for selectively binding the drug on the functionalized monomer within a weakly cross-linked polymeric hydrogel; forming the polymeric hydrogel comprising: forming a solution comprising amounts of bio-template, the functionalized monomer and a cross-linking monomer; complexing the functionalized monomer and the bio-template in the solution through non-covalent interactions, and initiating copolymerization of the functionalized monomer and the cross-linking monomer; forming the weakly cross-linked polymeric hydrogel into contact lens; and loading the contact lens with the drug, wherein the drug complexes with the functionalized monomer as a result of the monomer having the functional groups mimicking the receptor sites at the biological tissue wherein the drug is identified as having a biological mechanism of action.

While multiple embodiments are disclosed, still other embodiments of the present invention will become apparent to those skilled in the art from the following detailed description, which shows and describes illustrative embodiments of the invention. Accordingly, the drawings and detailed description are to be regarded as illustrative in nature and not restrictive.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a block diagram showing the steps for making contact lenses, in accordance with the embodiments of the invention.

FIGS. 2A-D illustrate the formation of a recognitive polymeric hydrogel, in accordance with the embodiments of the invention.

FIG. 3 illustrates a block diagram outlining steps for making functionalized monomer used in the synthesis of recognitive polymeric hydrogels, in accordance with the embodiments of the invention.

FIGS. 4A-C illustrate examples of sets of molecules that match, resemble or mimic each other.

FIGS. 5A-B are graphs that compare Ketotifen equilibrium isotherms in water for a recognitive polymeric hydrogel and a control hydrogel.

FIG. 5C graphs drug loading for recognitive polymeric hydrogels of the present invention against control hydrogels to show the enhanced drug loading for recognitive polymeric hydrogels of the present invention.

FIG. 6 shows a graph of drug release profiles for therapeutic contact lenses, in accordance with the embodiments of the invention.

FIG. 7A-B show graphs of drug release profiles for recognitive polymeric hydrogels, in accordance with the embodiments of the invention.

FIG. 8 is a graph that compares fluconazole equilibrium isotherms in water for a recognitive polymeric hydrogel and a control hydrogel.

FIG. 9 shows Fluconazole binding isotherm of 5% cross-linked poly-HEMA-co-AM-co-DEAEM-co-NVP-co-PEG200DMA hydrogel lenses (N=3, T=25 C). Percentage denotes percent mole cross linker per mole total monomers in feed.

FIG. 10 shows Fluconazole dynamic release in artificial lacrimal solution from 5% cross-linked poly-HEMA-co-AM-co-DEAEM-co-NVP-co-PEG200DMA hydrogel lenses (N=3, T=25 C). Percentage denotes percent mole cross linker per mole total monomers in feed.

FIG. 11A illustrates drug loading differences between free radial polymerization and living polymerization and iniferter reaction strategies.

FIG. 11B illustrates fractional release of diclofenac sodium template for poly(DEAEM-co-HEMA-co-PEG200DMA) gels.

FIG. 12 is a plot of the release profiles of HA of different sizes normalized to the maximum amount of HA released.

FIG. 13 shows the cumulative HA releases from lenses made with formulations containing 2 mg/g, 6.5 mg/g and 40 mg/g of HA in monomers/macromers (PM).

FIG. 14 highlights the release from only the 6.5 mg/g formulation.

FIG. 15 highlights the release from the basic formulation with no added monomers.

FIG. 16 highlights early studies including the functional monomers in the formulation at a 1% functional monomer.

FIG. 17 shows the release profiles of various formulations containing different weight percentages of functional monomers in the hydrogel.

FIG. 18 shows a graph of diffusion coefficients and orders of release for different lenses.

FIG. 19 shows the diffusion coefficient dependence on the polymer volume fractions.

FIG. 20 shows a plot of diffusion coefficients versus monomer percentage.

FIG. 21 highlights two formulations with 0.125% DEAEMA.

FIG. 22 shows the release of all diversified gels.

FIG. 23 shows the influence of the different functional monomers on the HA diffusion coefficient.

FIGS. 24-25 show how the polymer volume fraction varies with respect to various parameters.

FIGS. 26-27 show that the equilibrium weight and swelling ratios and polymer volume fraction in the swollen state of gels pre-HA release were equivalent to gels that were studied after HA release.

FIG. 28 highlights the optical transmission studies of swollen hydrogels.

FIG. 29 shows the percent change in concentration over time of a large HA.

FIG. 30 shows release profiles.

FIG. 31 illustrates a flowchart of a method of making a comfort agent delivery system.

FIG. 32 illustrates a flowchart of a method of dispensing a comfort agent in an eye biological tissue.

FIGS. 33A-B show microfluidic device schematics according to embodiments of the invention mimicking in-vivo drug release profiles in the eye.

FIGS. 34A-B show graphs of the equilibrium weight swelling ratio of therapeutic contact lenses, in accordance with the embodiments of the invention, (A) without additional cross linkers in comparison (B) with additional cross linkers.

FIG. 35 shows a graph of the dynamic release of HPMC from therapeutic contact lenses without imprinting and varying HPMC content of the lenses.

FIGS. 36A-B show graphs of the optical clarity (percentage transmittance) of therapeutic contact lenses according to embodiments of the invention in relationship with lenses prepared with additional cross linkers and/or imprinted lenses prepared with additional cross linkers.

FIG. 37 shows a graph of the drug release profiles for therapeutic contact lenses in accordance with the embodiments of the invention, showing the effect of the addition of cross linkers on release rate.

FIG. 38 shows a graph of the effect of monomer to bio-template ratio (M/T ratio) on the drug release profiles for therapeutic contact lenses in accordance with the embodiments of the invention.

FIG. 39 shows a graph of the effect of extended release profiles from imprinted therapeutic contact lenses in accordance with the embodiments of the invention.

FIG. 40 shows a graph of the optical clarity of imprinted therapeutic contact lenses in accordance with the embodiments of the invention.

FIG. 41 shows a graph of the dynamic release in physiological flow rate of imprinted therapeutic contact lenses in accordance with the embodiments of the invention.

Various embodiments of the present invention will be described in detail with reference to the drawings, wherein like reference numerals represent like parts throughout the several views. Reference to various embodiments does not limit the scope of the invention. Figures represented herein are not limitations to the various embodiments according to the invention and are presented for exemplary illustration of the invention.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENT

Hydrogels are insoluble, cross-linked polymer network structures composed of hydrophilic homo- or hetero-co-polymers, which have the ability to absorb significant amounts of water. Consequently, this is an essential property to achieve an immunotolerant surface and matrix (i.e., with respect to protein adsorption or cell adhesion). Due to their significant water content, hydrogels also possess a degree of flexibility very similar to natural tissue, which minimizes potential irritation to surrounding membranes and tissues.

The hydrophilic and hydrophobic balance of a gel carrier can be altered to provide tunable contributions that present different solvent diffusion characteristics, which in turn influence the diffusive release of a drug contained within the gel matrix. In general, one may polymerize a hydrophilic monomer with other less hydrophilic or more hydrophobic monomers to achieve desired swelling properties.

These techniques have led to a wide range of swellable hydrogels. Knowledge of the swelling characteristics is of major importance in biomedical and pharmaceutical applications since the equilibrium degree of swelling influences the diffusion coefficient through the hydrogel, surface properties and surface mobility, mechanical properties, and optical properties. Drug release depends on two simultaneous rate processes: water migration into the network and drug diffusion outward through the swollen gel.

Soft contact lenses are made of hydrogels. The typical material properties for contact lenses involve a number of considerations such as optical quality (good transmission of visible light), high chemical and mechanical stability, manufacturability at reasonable cost, high oxygen transmissibility, tear film wettability for comfort, and resistance to accumulation of protein and lipid deposits, as well as a suitable cleaning and disinfecting scheme.

Soft contact lenses typically consist of poly(2-hydroxyethyl methacrylate) (PHEMA). Other lens materials include HEMA copolymerized with other monomers such as methacrylic acid, acetone acrylamide, and vinyl pyrrolidone. Also, commonly used are copolymers of vinyl pyrrolidone and methyl methacrylate as well as copolymers of glycerol methacrylate and methyl methacrylate. Minor ingredients have included a variety of other monomers as well as cross-linking agents.

The immersion and soaking of soft contact lenses in drug solutions has shown promise in the increase of drug bioavailability with a minimization of side effects. However, the materials and constituent chemistry of the macromolecular chains and subsequent interaction with drugs is random and typically leads to poor drug loading.

In order to address the above referenced shortcomings, the present invention is directed to the use of biomimetic imprinting of hydrogels to make hydrogels matrices that can selectively bind a drug through complexing sites leading to improved loading of a drug and controlled time release of the drug. These hydrogels are referred to as recognitive polymeric hydrogels. The polymerization reaction forms the contact lenses, which can be used to administer drugs through contact with the eyes, thereby replacing traditional eye drop therapies. Alternatively, the recognitive polymeric hydrogels can be formed or fashioned into contact lenses which can be used to administer drugs through contact with the eyes, thereby replacing traditional eye drop therapies or other mechanisms of delivery.

For example, ketotifen fumarate is a potent fast acting and highly selective histamine H1 antagonists with a sustained duration of action. Ketotifen fumarate inhibits itching, redness, eyelid swelling, tearing, and chemosis induced by conjunctival provocation with allergens and histamine. With topical application in the form of eye drops, absorption is incomplete and bioavailability is low. Thus, the dose is usually administered multiple times daily. Also, due to a high concentration of drug and other constituents of the ophthalmic suspension preparation, patients are advised not to wear soft contact lenses. Accordingly, a soft contact lens that could be used to administer ketotifen fumarate would not only enhance the efficacy of the treatment, but also allow allergy sufferers to wear contact lenses.

Referring to FIG. 1 which is a block diagram 100 outlining steps for making contact lenses, in accordance with the embodiments of the invention and FIG. 2 which is a graphical representation of forming a recognitive polymeric hydrogel 221. In the step 101, the recognitive hydrogel matrix 221 is formed. The recognitive hydrogel 221 is formed by generating a solution 200 comprising one or more bio-template 201, one or more functionalized monomers 203 and 203′, one or more cross-linking monomers 205 with or without a solvent. In the solution 200′ the functionalized monomers 203 and 203′ complexes with the bio-templates 201. A suitable initiator or mixture initiators 207 is used to co-polymerize the functionalized monomers 203 and 203′ with a cross-linking monomer 205 to form the loaded hydrogel 220 comprising a hydrogel matrix 221 with bio-templates 201 complexing at site 209 through the hydrogel matrix 221.

Preferably, the bio-templates are complexed with the hydrogel matrix 221 through weak or non-covalent interactions, as explained above, whereby the bio-templates can be washed or rinsed from the complexed hydrogel 220 to form an un-complexed recognitive polymeric hydrogel 221, which has vacant complexing sites 209 that can be used to complex drug molecules that are structurally and/or chemically similar to the bio-templates 201. It will be clear from the discussions above and below that the bio-templates can be a drug and, therefore, washing the bio-templates from the hydrogel matrix 221 may not be necessary for all drug delivery systems that are synthesized.

Still referring to both FIG. 1 and FIG. 2, after the recognitive hydrogel 221 is formed, in the step 101, in the step 103 the recognitive hydrogel 221 can be formed into contact lenses using any technique known in the art. It is understood that the step the step 103 is not necessary, when the polymerization reaction forms the contact lenses, such as described previously. Where the bio-template is a drug, the contact lenses can be placed in contact with eyes in the step 107 to administer or deliver the drug to or through the eyes. Where, the bio-template 201 has been washed from the recognitive hydrogel matrix prior to or after the step 103 of forming the contact lenses from the recognitive hydrogel matrix, then in the step 109 or the step 105, respectively, the recognitive hydrogel matrix or the contact lenses are loaded with a drug. The recognitive hydrogel matrix or the contact lenses can be loaded with the drug by soaking the recognitive hydrogel matrix or the contact lenses in an aqueous drug solution.

Now referring to FIG. 2 and FIG. 3. In accordance with further embodiments of the invention prior to the step of making an ophthalmic drug delivery system, such as described with reference to FIG. 1, in the step 301 the target tissue to be treated with the drug or biological mechanism of action is studied to determine the types of molecules or functional groups that are associated with the action of the drug at the target tissue to effect the target tissue. Based on this information, in the step 303, functionalized monomers are synthesized with functional groups that mimic or resemble molecules or functional groups that are associated with the action of the drug at the target tissue. The functionalized monomers with the functional groups that mimic or resemble molecules or functional groups that are associated with metabolizing the drug at the target tissue are then used to synthesize a drug delivery system, such as described above with reference to FIG. 1. The biomimetic approach is the processes of mimicking biological recognition or exploiting biological mechanisms. Specifically, it is the process of coordinating biological molecular recognition, interactions, or actions to design materials that can be structurally similar to and/or function in similar ways as biological structures.

FIGS. 4A-C illustrate examples of sets of molecules that match, resemble or mimic each other. With reference to the bio-mimetic approach for synthesizing recognitive hydrogel polymers described above, acrylic acid can be used to mimic aspartic acid (FIG. 4A), acrylaminde can be used to mimic asparagine (FIG. 4B) and N-vinyl pyrrolidinone can be used to mimic tyrosine (FIG. 4C). Aspartic acid, asparagine, and tyrosine are known to be of the group of amino acids providing the non-covalent interactions in the ligand binding pocket for histamine. For example, structural analysis of ligand binding pockets and amino acids involved in multiple non-covalent binding points provide one of many rational frameworks to synthesize recognitive networks from functional monomers. Antihistamine has been shown to bind more tightly and have a higher affinity than histamine for the histamine binding pocket.

Compositions

In accordance with the embodiments of the invention a hydrogel matrix that is formed from silicone-based cross-linking monomers, carbon based or organic-based monomers, macromers or a combination thereof. Suitable cross-linking monomers include but are not limited to Polyethylene glycol (200) dimethacrylate (PEG200DMA), ethylene glycol dimethacrylate (EGDMA), tetraethyleneglycol dimethacrylate (TEGDMA), N,N′-Methylene-bis-acrylamide and polyethylene glycol (600) dimethacrylate (PEG600DMA). Suitable silicone-based cross-linking monomers can include tris(trimethylsiloxy)silyl propyl methacrylate (TRIS) and hydrophilic TRIS derivatives such as tris(trimethylsiloxy)silyl propyl vinyl carbamate (TPVC), tris(trimethylsiloxy)silyl propyl glycerol methacrylate (SIGMA), tris(trimethylsiloxy)silyl propyl methacryloxyethylcarbamate (TSMC); polydimethylsiloxane (PDMS) and PDMS derivatives, such as methacrylate end-capped fluoro-grafted PDMS cross linker, a methacrylate end-capped urethane-siloxane copolymer cross linker, a styrene-capped siloxane polymer containing polyethylene oxide and polypropylene oxide blocks; and siloxanes containing hydrophilic grafts or amino acid residue grafts, and siloxanes containing hydrophilic blocks or containing amino acid residue grafts. The molecular structure of these monomers can be altered chemically to contain moieties that match amino acid residues or other biological molecules. In cases where the above monomers, when polymerized with hydrophilic monomers, a solubilizing cosolvent may be used such as dimethylsulfoxide (DMSO), isopropanol, ethanol, etc. or a protecting/deprotecting group strategy.

Crosslinking monomer(s) or macromer(s) amounts can be from (0.1 to 40%, moles crosslinking monomer/moles all monomers); Functional monomers, 99.9% to 60% (moles functional monomer/moles all monomers) with varying relative portions of multiple functional monomers; initiator concentration ranging from 0.1 to 30 wt %; solvent concentration ranging from 0% to 50 wt % (but no solvent is preferred); monomer to bio-template ratio (M/T) ranging from 0.1 to 5,000, preferably 200 to 1,000, with 950 preferred for the ketotifen polymers presented herein, under an nitrogen or air environment (in air, the wt % of initiator should be increased above 10 wt %.

As one skilled in the art will appreciate, the size of monomers and/or macromers (e.g. long chain monomer) can also be adjusted with the extent of degree of substitution of double bonds on the polymer chains. Such monomers and/or macromers may be included within the crosslinking monomers or macromers of the present invention. In such aspects these alterations to the monomers and/or macromers may result in the increase in crosslinking amount according to the hydrogel delivery systems of the invention. In an aspect, the crosslinking monomer(s) and/or macromer(s) amounts can be up to at least about 80% (moles crosslinking monomers and/or macromer/moles all monomers and/or macromers). In a further aspect, the crosslinking monomer(s) and/or macromer(s) amounts can be from 0.1% to 80%, from 0.1% to 70%, from 0.1% to 60%, from 0.1% to 50%, or from 0.1% to 0% (moles crosslinking monomers and/or macromer/moles all monomers and/or macromers). Without limiting the scope of the invention, all amounts within these cited ranges are included within the embodiments of the invention.

In a further aspect, the changes to the crosslinking and/or crosslinking monomers and/macromers according to the invention will result in functional alterations to the hydrogel contact lenses. In an aspect, variations in the extent of crosslinking according to the invention can result in two spectrums of modifications to the contact lenses. For example, in the event the crosslinking monomers and/or macromers provide a structure that is too flexible, the bioimprinting methods will not function due to the crosslinking chains of the monomers and/or macromers being too far apart (e.g. distanced and/or spread apart). As a result, the multiple chains will not interact with the bio-template (including drug and/or comfort agent) to form the contact lenses according to the invention and the biomimetic approach. Alternatively, in the event the extent of crosslinking is too rigid, the hydrogel contact lenses will not function as an extended wear lens and/or a comfortable soft contact lens. Therefore, the compositions and methods according to the invention provide crosslinking monomers and/or macromers that provide sufficient rigidity to allow the formation of the bio-template using the bioimprinting methods, which not forming too rigid of a contact lenses which will be uncomfortable on the eye and largely not suitable for extended wear and delivery of a drug and/or comfort agent. The ophthalmic drug delivery system also includes a bio-template, that is drug molecules, prodrugs, protein, amino acid, proteinic drug, oligopeptide, polypeptide, oligonucleotide, ribonucleic acid, deoxyribonucleic acid, antibody, vitamin, or other biologically active compound. This also includes a drug with an attached bio-template. The bio-template is preferably bound to the hydrogel matrix through one or more of electrostatic interactions, hydrogen bonding, hydrophobic interactions, coordination complexation, and Van der Waals forces.

Bio-templates are preferably weakly bound to a hydrogel matrix through functionalized monomer units, macromer units or oligomer units that are co-polymerized into the hydrogel matrix to form receptor locations within the hydrogel matrix that resemble or mimic the receptor sites or molecules associated with the biological target tissue to be treated with the drug or the biological mechanism of action

In accordance with the embodiments of the invention, a portion of the bio-template can be washed out from the recognitive hydrogel polymer, loaded with a drug. The polymerization reaction forms a contact lens. For example, the gel is polymerized in a mold or compression casting. After contact lenses are formed they can be used to administer the drug through contact with eyes. Alternatively, the recognitive hydrogel polymer can be formed into contact lenses, washed to remove a portion of the bio-template and then loaded with the drug. Where the bio-template is the drug, the washing step can be illuminated or truncated. In formulations where the bio-template is a drug, the free base form of the drug or hydrochloride salt of the drug can be used.

In accordance with the method of the present invention, a biomimetic recognitive polymeric hydrogel is formed by making a mixture or solution that includes amounts of a bio-template or drug, functionalized monomer or monomers, cross-linking monomer or monomers and polymerization initiator in a suitable solvent or without solvent. Suitable initiators include water and non-water soluble initiators, but are not limited to azobisisobutyronitrile (AIBN), 2,2-dimethoxy-2-phenyl acetophenone (DMPA), 1-hydroxycyclohexyl phenyl ketone (Irgacure® 184), 2,2-dimethoxy-1,2-diphenylethan-1-one (Irgacure 651), ammonium persulfate, iniferter such as tetraethylthiuram disulfide, or combinations thereof. The polymerization can be photo-initiated, thermally-initiated, redox-initiated or a combinations thereof.

The functionalized monomer or monomers complex with the bio-template and copolymerize with cross-linking monomer or monomers to form a biomimetic recognitive polymeric hydrogel, such as described above. Functional or reactive monomers useful herein are those which possess chemical or thermodynamic compatibility with a desired bio-template. As used herein, the term functional monomer includes moieties or chemical compounds in which there is at least one double bond group that can be incorporated into a growing polymer chain by chemical reaction and one end that has functionality that will interact with the bio-template through one or more of electrostatic interactions, hydrogen bonding, hydrophobic interactions, coordination complexation, and Van der Waals forces. Functional monomers include macromers, oligomers, and polymer chains with pendent functionality and which have the capability of being cross-linked to create the recognitive hydrogel. Crosslinking monomers include macromers, oligomers, and includes chemicals with multiple double bond functionality that can be polymerized into a polymer network. Crosslinking can also include covalent or non-covalent bonding of chemistry of differing polymer chains, including for example crosslinking without a double bond (including a covalent reaction with a group along the polymer chain).

Initiator-chain transfer molecules, iniferters, have been used to produce well controlled block copolymers, polymers of low polydispersity, and graft polymers as well as cross-linked polymer systems on surfaces. Of utmost importance is the control over the polymerization and associated network structure, which depends on the dynamic equilibrium between active and dormant species. Conventional free-radical polymerization is highly non-ideal and differences in theory and experimental data indicate heterogeneity within the network structure.

In order to control the polymerization reaction further by altering the kinetic chain length and potentially increasing the number of recognition sites and the quality of recognition (e.g., alter the structural architecture of the polymeric network to affect template drug capacity, affinity, and diffusional transport), we investigated the use of initiator-chain transfer molecules, iniferters. By using the iniferter, tetraethylthiuram disulfide (TED), the number of binding sites was dramatically increased at approximately equivalent binding affinity; however both affinity and capacity can be increased. Iniferters used in this work decay into two dithiocarbamyl radicals (DTC); which are more stable compared to carbon radicals. The stability of the iniferter produced radical negates its significance on the initiation and propagation steps during the polymerization reaction, which in this particular case required the addition of carbon radicals, AIBN, to initiate the polymerization reaction. During termination steps of the polymerization reaction, the stable DTC radicals reversibly terminate with growing polymer radical chains which forms a chain that can re-absorb UV light and decay back into a polymer radical and a DTC radical. The limitations and structural heterogeneity of radical polymerizations caused by fast termination reactions can be reduced since iniferters provide a reversible termination reaction.

Examples of living or controlled polymerization include, but are not limited to living anioinic or cationic polymerization, ring opening metathesis polymerization (ROMP), group transfer polymerization (GOP), living Ziegler-Natta polymerization, and free-radical polymerization (e.g., iniferter polymerization, catalytic chain transfer polymerizaton, stable free radical mediated polymerization (SFRP), ATRF or atom transfer radical polymerization, reversible addition fragmentation chain transfer (RAFT) polymerization, Iodine Transfer polymerization, Selenium-centered mediated polymerization, Telluride-mediated polymerization (TERP), Stibine-mediated polymerization).

Examples of functionalized monomers include, but are not limited to, 2-hydroxyethylmethacrylate (HEMA), Acrylic Acid (AA), Acrylamide (AM), N-vinyl 2-pyrrolidone (NVP), 1-vinyl-2-pyrrolidone (VP), methyl methacrylate (MMA), methacrylic acid (MAA), acetone acrylamide, 2-ethyl-2-(hydroxymethyl)-1,3-propanediol trimethacrylate, N-(1,1-dimethyl-3-oxobutyl)acrylamide, 2-ethyl-2-(hydroxymethyl)-1,3-propanediol trimethacrylate, 2,3-dihydroxypropyl methacrylate, allyl methacrylate, 3-[3,3,5,5,5-pentamethyl-1,1-bis[pentamethyldisiloxanyl)oxy]trisiloxanyl]propyl methacrylate, 3-[3,3,3-trimethyl-1,1-bis(trimethylsiloxy)disiloxanyl]propyl methacrylate (TRIS), N-(1,1-dimethyl-3-oxybutyl)acrylamide, dimethyl itaconate, 2,2,2,-trifluoro-1-(trifluoromethyl)ethyl methacrylate, 2,2,2-trifluoroethyl methacrylate, methacryloxypropylbis(trimethylsiloxy)methylsilane, methacryloxypropylpentamethyldisiloxane, (3-methacryloxy-2-hydroxypropyloxy)propylbis(trimethylsiloxy)methylsilane, 4-t-butyl-2-hydroxycyclohexyl methacrylate, dimethylacrylamide and glycerol methacrylate. Once formed the biomimetic recognitive polymeric hydrogel can be formed into contact lenses or as described above the polymerization reaction forms the contact lenses.

In accordance with further embodiments of the invention, functionalized monomers are synthesized or selected by identifying receptor sites or molecules associated with the target biological tissue to be treated by the drug or that are associated with metabolizing the drug. Then functionalized portions of the functionalized monomers are synthesized to chemically and/or structurally resemble or mimic the receptor sites or molecules that are associated with the biological mechanism of action of the drug. These functionalized monomers are then copolymerized with the cross-linking monomer or monomers used to form the hydrogel matrix, such as described above.

After the drug has been depleted from the contact lenses through the eyes, the contact lenses can be re-loaded with the drug by soaking the contact lenses in the reconstituting drug solution. While the contact lens have been described in detail as being used to deliver antihistamines and other allergy drugs, ophthalmic drug delivery systems and methods of the present invention can be used to deliver any number of drugs through contact on the eye and/or systemically.

Drugs that can be delivered by the system and method of the present invention include, but are not limited to, Anti-bacterials Anti-infectives and Anti-microbial Agents (genteelly referred to as antibiotics) such as Penicillins (including Aminopenicillins and/or penicillinas in conjunction with penicillinase inhibitor), Cephalosporins (and the closely related cephamycins and carbapenems), Fluoroquinolones, Tetracyclines, Macrolides, Aminoglycosides. Specific examples include, but are not limited to, erythromycin, bacitracin zinc, polymyxin, polymyxin B sulfates, neomycin, gentamycin, tobramycin, gramicidin, ciprofloxacin, trimethoprim, ofloxacin, levofloxacin, gatifloxacin, moxifloxacin, norfloxacin, sodium sulfacetamide, chloramphenicol, tetracycline, azithromycin, clarithyromycin, trimethoprim sulfate and bacitracin.

The ophthalmic drug delivery system and method of the present invention can also be used to deliver Non-steroidal (NSAIDs) and Steroidal Anti-inflammatory Agents (genteelly referred to as anti-inflammatory agents) including both COX-1 and COX-2 inhibitors. Examples include, but are not limited to, corticosteroids, medrysone, prednisolone, prednisolone acetate, prednisolone sodium phosphate, fluormetholone, dexamethasone, dexamethasone sodium phosphate, betamethasone, fluoromethasone, antazoline, fluorometholone acetate, rimexolone, loteprednol etabonate, diclofenac (diclofenac sodium), ketorolac, ketorolac tromethamine, hydrocortisone, bromfenac, flurbiprofen, antazoline and xylometazoline.

The ophthalmic drug delivery system and method of the present invention can also be used to deliver Anti-histamines, Mast cell stabilizers, and Anti-allergy Agents (generally referred to as anti-histamines). Examples include, but are not limited, cromolyn sodium, lodoxamide tromethamine, olopatadine HCl, nedocromil sodium, ketotifen fumarate, levocabastine HCL, azelastine HCL, pemirolast (pemirolast potassium), epinastine HCL, naphazoline HCL, emedastine, antazoline, pheniramine, sodium cromoglycate, N-acetyl-aspartyl glutamic acid and amlexanox.

In yet further embodiments of the invention the ophthalmic drug delivery system and method are used to deliver Anti-viral Agents including, but not limited to, trifluridine and vidarabine; Anti-Cancer Therapeutics including, but not limited to, dexamethasone and 5-fluorouracil (5FU); Local Anesthetics including, but are not limited to, tetracaine, proparacaine HCL and benoxinate HCL; Cycloplegics and Mydriatics including, but not limited to, Atropine sulfate, phenylephrine HCL, Cyclopentolate HCL, scopolamine HBr, homatropine HBr, tropicamide and hydroxyamphetamine Hbr; Comfort Molecules or Molecules (generally referred as lubricating agents) to treat Keratoconjunctivitis Sicca (Dry Eye) including, but not limited to, Hyaluronic acid or hyaluronan (of varying Molecular Weight, MW), hydroxypropyl cellulose (of varying MW), hydroxypropylmethyl cellulose (of varying MW), gefarnate, hydroxyeicosatetranenoic acid (15-(S)-HETE), phospholipid-HETE derivatives, phoshoroylcholine or other polar lipids, carboxymethyl cellulose (of varying MW), polyethylene glycol (of varying MW), polyvinyl alcohol (of varying MW), rebamipide, pimecrolimus, ecabet sodium and hydrophilic polymers; Immuno-suppressive and Immuno-modulating Agents including, but not limited to, Cyclosporine, tacrolimus, anti-IgE and cytokine antagonists; and Anti-Glaucoma Agents including beta blockers, pilocarpine, direct-acting miotics, prostagladins, alpha adrenergic agonists, carbonic anhydrase inhibitors including, but not limited to betaxolol HCL, levobunolol HCL, metipranolol HCL, timolol maleate or hemihydrate, carteolol HCL, carbachol, pilocarpine HCL, latanoprost, bimatoprost, travoprost, brimonidine tartrate, apraclonidine HCL, brinzolamide and dorzolamide HCL; decongestants, vasodilaters vasoconstrictors including, but not limited to epinephrine and pseudoephedrine.

In yet further embodiments of the invention the ophthalmic drug delivery system and method are used to deliver anti-microbial agents including, but not limited to, anti-fungal agents. A number of different classes of anti-microbial or anti-fungal agents are considered to be within the scope of the present invention. Suitable classes of anti-microbial or anti-fungal agents include, but are not limited to, Allyamines, other non-azole ergosterol biosynthesis inhibitors, antimetabolities, Azoles, Chitin Synthase Inhibitors, Glucan Synthesis Inhibitors, and Polyenes.

Examples of Allyamines include, but are not limited to, Amorolfine, Butenafine, Naftifine and Terbinafine; an example of an antimetabolite includes, but is not limited to, Flucytosine; examples of Azoles include, but are not limited to, Imadazoles and Triazoles, such as Fluconazole, Itraconazole, Ketoconazole, Posaconazole, Ravuconazole, Voriconazole, Clotrimazole, Econazole, Miconazole, Miconazole Nitrate, Oxiconazole, Sulconazole, Terconazole, Tioconazole and Enilconazole; examples of Chitin Synthase Inhibitors include, but are not limited to, Nikkomycin Z and Lufenuron; examples of Echinocandins or Glucan Synthesis Inhibitors include, but are not limited to, Caspofungin, Micafugin and Anidulafungin, examples of Polyenes include, but are not limited to, Amphotericin B (AmB), Natamycin, Pimaricinand Nystatin. Other suitable anti-microbial or anti-fungal agents include, but are not limited to, Chlorhexidine gluconate, Griseofulvin, Ciclopirox Olamine, Haloprogin, Tolnaftate, Undecylenate, Povidone iodine, Silver sulfadiazine and antimicrobial peptides and glycoproteins (e.g., such as lactoferrin, ambicin, nisin, polymixin B, gramicidin S, etc.).

An ophthalmic drug delivery system of the present invention has particular applications for treating animals, including large animals such as horses diagnosed with equine fungal keratitis with an anti-microbial agents or anti-fungal agents such as those listed above. Equine fungal keratitis is a corneal disease that accounts for 13% of all equine corneal problems. It is generally caused by corneal trauma, usually associated with injury by organic matter. The injury allows fungal microorganisms to invade the cornea and cause ulcers, anterior uvetis, corneal perforation and iris prolapse. The typical treatment for equine fungal keratitis is frequent, expensive and prolonged treatment. This is due to the number of ocular treatment barriers and the difficulty of dealing with large animals.

For example, a single drop of topical drop (the usual topical dose) has a residence time of between 2-23 minutes. As a result, the bioavailability of the drug is approximately 10% of the administered dose. About 44% of animals diagnosed with equine fungal keratitis go blind and a number of these cases require that the eye is removed to prevent further spreading of infection. The physiological properties of the eye hinder the effectiveness of topical ocular drugs. Because of the effectiveness of the tear film in flushing out foreign substances, the goal of this novel ocular drug delivery method is to increase the bioavailability by releasing a smaller concentration of the drug at a constant rate, thereby increasing the corneal contact time.

EXAMPLE 1

Materials and Methods: Acrylic Acid (AA), Acrylamide (AM), N-Vinyl-2-Pyrrolidone (NVP) and 2-hydroxyethylmethacrylate (HEMA), Azobisisobutyronitrile (AIBN), and Ketotifen Fumarate were purchased from Sigma-Aldrich. Polyethylene glycol (200) dimethacrylate (PEG200DMA) was purchased from Polysciences, Inc. All chemicals were used as received. Polymer and copolymer networks were made using various mixtures of above monomers (e.g. Poly(AA-co-AM-HEMA-PEG200DMA), Poly (AA-co-HEMA-co-PEG200DMA), Poly (AM-co-HEMA-co-PEG200DMA), Poly(AA-co-AM-co-NVP-co-HEMA-PEG200DMA)). Current work is directed to producing networks that can also be used in the formation of contact lens for anti-histamines with monomers and copolymers of molecules such as N-vinyl 2-pyrrolidone (NVP), 1-vinyl-2-pyrrolidone (VP), methyl methacrylate (MMA), methacrylic acid (MAA), acetone acrylamide, ethylene glycol dimethacrylate (EGDMA), 2-ethyl-2-(hydroxymethyl)-1,3-propanediol trimethacrylate, N-(1,1-dimethyl-3-oxobutyl)acrylamide, 2-ethyl-2-(hydroxymethyl)-1,3-propanediol trimethacrylate, 2,3-dihydroxypropyl methacrylate, allyl methacrylate any other suitable monomers, such as those referenced previously.

Accurate quantities of monomers, template molecules and cross linkers were added in that order, and the mixture was sonicated to obtain a homogenous solution. In particular, a typical formulation consisted of 5 mole % cross-linking monomer (PEG200DMA) in a solution of Acrylamide (M), HEMA (M), Ketotifen (T), with an M/T ratio of approximately 950 (92% HEMA, 1% of remaining monomers, and approximately 1 mole % drug depending on the M/T ratio). Controls were also prepared without the template. Next, initiator AIBN was added in low light conditions, and the solutions were allowed to equilibrate for 12 hours in darkness. This step allowed the monomers and template to orient themselves and reach their free energy minima, thus beginning the configurational imprinting at the molecular level. However, this step occurs very quickly such as on the order of minutes.

The solutions were then transferred to an MBRAUN Labmaster 130 1500/1000 Glovebox, which provides an inert nitrogenous and temperature-controlled atmosphere for free-radical photopolymerization. With an increase in photoinitiator wt. %, this step can proceed in air. The solutions were uncapped and left open to the nitrogen until the oxygen levels reached negligible levels (<0.1 ppm). The solutions were inserted into glass molds (6 in. by 6 in.) separated by a Teflon frame 0.8 mm wide, as measured by a Vernier caliper. The glass plates were coated with chlorotrimethylsilane in order to prevent the polymer matrix from sticking to the glass, as it demonstrates a strong adherent tendency due to hydrogen bonding. Polymerization was carried out for ten minutes at 325 V using a Dymax UV light source. The intensity of radiation was 40 mW/cm², as measured with a radiometer, and the temperature was 36° C., as measured by a thermocouple.

The polymer was peeled off the glass plates with flowing deionized water (Millipore, 18.2 mO·cm, pH 6), and then was allowed to soften for approximately 10 minutes. Circular discs were cut using a Size 10 cork borer (13.5 mm), and were typically washed for 5 days in a continuous flow system using deionized water. All washes proceeded until the absence of detectable drug was verified by spectroscopic monitoring. To obtain dry weights, some discs were allowed to dry under laboratory conditions (20° C.) for 36 hours. The discs were then transferred to a vacuum oven (27 in. Hg, 33-34° C.) for 48 hours until they were dry (less than 0.1 wt % difference).

Polymer penetrant uptake and swelling data were obtained in deionized water with samples taken every 5 min. for the first hour, and then every hour for 10 hours until equilibrium was reached. As the gel was removed from the water, excess surface water was dabbed with a dry Kim wipe. The equilibrium weight swelling ratio at time t, q, for a given gel was calculated using the weights of the gels at a time and the dry polymer weights, respectively, using equations based on Archimedes principle of buoyancy. Dynamic and Equilibrium Template Binding: Dynamic template drug molecule binding was performed until equilibrium had been established for each system. Stock solutions of drug with a concentration 2 mg/ml were prepared and diluted with deionized water to produce 0.1, 0.2, 0.3, 0.4 and 0.5 mg/ml solutions. Each solution was vortexed for 30 seconds to provide homogeneity, and initial UV absorbances were noted. Gels were then inserted into the vials and were placed on a Stovall Belly Button Orbital Shaker over the entire duration of the binding cycle to provide adequate mixing. A 200 μL aliquot of each sample was placed in a Corning Costar UV-transparent microplate, and absorbance readings were taken using a Biotek Spectrophotometer at 268 nm. After measurement, the reading sample was returned to the original samples, to avoid fluctuations in concentrations due to sampling methods.

Dynamic Release Studies: In obtaining the preliminary results, dynamic release studies were conducted in DI water, artificial lacrimal fluid (6.78 g/L NaCl, 2.18 g/L NaHCO₃, 1.38 g/L KCl, 0.084 g/L CaCl₂.2H₂O, pH 8), and lysozyme (1 mg/ml) in artificial lacrimal fluid. Gels which had been drug loaded were placed in 30 ml of DI water, and the solutions were continuously agitated with a Servodyne mixer (Cole Palmer Instrument Co.) at 120 rpm. Release of drug was monitored at 268 nm by drawing 200 μL of solution into a 96-well Corning Costar UV-transparent microplate, and measurements were taken in a Synergy UV-Vis Spectrophotometer (Biotek). Absorbances were recorded for three samples, averaged, and corrected by subtracting the relevant controls. Solutions were replaced after each reading. Separate studies were conducted to determine if infinite sink conditions existed and those conditions were matched throughout all experiments.

Polymerization Kinetics and Network Formation: Solutions were prepared with 0, 0.1, 0.5, and 1 mole percent of Ketotifen in the initial monomer solutions. Kinetic studies were conducted with a differential scanning photocalorimeter (DPC, Model No. DSC Q100, TA Instruments with Mercury light source). Samples of 10 μL were placed in an aluminum hermetic pan and purged with nitrogen (flow rate 40 ml/min) in order to prevent oxidative inhibition. They were allowed to equilibrate at 35° C. for 15 minutes, before shining UV light at 40 mW/cm2 for 12 minutes.

The heat that evolved was measured as a function of time, and the theoretical enthalpy of the monomer solution was used to calculate the rate of polymerization, Rp, in units of fractional double bond conversion per second. Integration of the rate of polymerization curve versus time yielded the conversion as a function of time or reaction rate. The presence of template and a solvent, if used, was accounted for in the calculations, as it did not participate in the polymerization reaction. Experimental results were reproducible and the greatest source of error involved the assumed theoretical enthalpies in the calculations of the rate of polymerization and conversion. For all studies, the enthalpies were assumed to have errors of +5%. The assumptions in the copolymerization of two monomers (i.e., functional and cross-linking monomers) were that each monomer had equal reactivity and the theoretical enthalpy derived for a co-monomer mixture was an average of the enthalpies of individual monomers. The theoretical enthalpy of methacrylate double bonds was equal to 13.1 kcal mole-1 and the theoretical enthalpy of acrylate double bonds was equal to 20.6 kcal mole-1.

Results

FIG. 5A shows a graph 500 of the equilibrium binding isotherm for Ketotifen in water for Poly(acrylamide-co-HEMA-co-poly(ethylene glycol)200 dimethacrylate) hydrogel networks with a cross-linking percentage of 5%. N=3 and T=25° C. The recognitive hydrogel network is represented by the line 501 and the control hydrogel network is represented by the line 503. Percentage denotes percent mole cross linker per mole total monomers in feed.

FIG. 5B shows a graph 510 of the equilibrium binding isotherm for Ketotifen in water for Poly(acrylic acid-co-HEMA-co-poly(ethylene glycol)200 dimethacrylate) hydrogel networks with a cross-linking percentage of 5%. N=3 and T=25° C. The recognitive hydrogel networks is represented by line 511 and the control hydrogel network is represented by line 513. Percentage denotes percent mole cross linker per mole total monomers in feed.

FIG. 5C shows a graph 540 of enhanced Loading of Ketotifen for Multiple Monomer Gels for Poly(n-co-HEMA-co-poly(ethylene glycol)200 dimethacrylate) Networks. The Functional monomers uses are acrylic acid, acrylamide, NVP, or an equal mole mixture of both. The Recognitive networks are shown as hatched bars 543 and the Control networks are shown as clear bars 541.

FIG. 6 shows a graph 600 of Tailorable Release Profiles Of Therapeutic Contact Lenses for Poly(n-co-HEMA-co-poly(ethylene glycol)200 dimethacrylate) Networks in Artificial Lacrimal Fluid, where n is AM (represented by circles), AA (represented by squares), AA-AM(represented by triangles), and NVP-AA-AM (represented by diamonds) recognitive networks respectively. Results demonstrate approximately constant release rate of ketotifen fumarate for 1 to 5 days.

FIG. 7A shows a graph 700 of Release Data for Poly(AM-co-HEMA-co-poly(ethyleneglycol)200 dimethacrylate) Recognitive Networks. Fraction of Mass Released in Artificial Lacrimal Solution With/Without Lysozyme.

FIG. 7B shows a graph 725 of Release Data for Poly(AM-co-AA-co-HEMA-co-poly(ethyleneglycol)200 dimethacrylate) Networks Mass of Drug Released in Artificial Lacrimal Solution.

FIG. 8 shows a graph 800 of the equilibrium binding isotherm for fluconazole in water for Poly(acrylic acid-co-HEMA-co-poly(ethylene glycol)200 dimethacrylate) hydrogel networks with a cross-linking percentage of 5%. N=3 and T=25° C. The recognitive hydrogel network is represented by the line 811 and the control hydrogel network is represented by the line 813. Percentage denotes percent mole cross linker per mole total monomers in feed.

I. Enhanced Loading and Performance of Multiple Monomer Mixtures

In the preliminary work, hydrogels were produced with enhanced loading for ketotifen fumarate. Polymers were made with the following monomers: acrylic acid (AA), N-vinyl 2-pyrrolidone (NVP), acrylamide (AM), 2-hydroxyethylmethacrylate (HEMA), and polyethylene glycol (200) dimethacrylate (PEG200DMA).

We hypothesized that gels composed of multiple functional monomers would outperform those composed of single functional monomers. For anti-histamine recognitive polymers, this would better mimic the docking site of histamine at the molecular level providing all the relevant functionality necessary for non-covalent interactions. We have proved that loading properties of gels are improved with multiple monomer mixtures.

Gels of multiple complexation points with varying functionalities outperformed the gels formed with less diverse functional monomer and showed the highest maximum bound of ketotifen and highest difference over control gels. Equilibrium binding isotherms for Poly(AM-co-AA-co-HEMA-co-PEG200DMA) networks demonstrate enhanced loading with a factor of 2 times increase in the loading of drug compared to conventional networks (i.e., gels prepared without template and comparable to existing contact lenses) depending on polymer formulation and polymerization conditions. Poly(AM-co-HEMA-co-PEG200DMA) networks demonstrated a factor of 2 or 100% increase in the loading of drug compared to control networks with lower bound amounts. Poly(AA-co-HEMA-co-PEG200DMA) networks show a factor of 6 times increase over control in the loading of ketotifen with the overall drug bound being the lowest of the polymer formulations studies (approximately 33% less ketotifen loading than the AM functionalized network).

For all systems, an increase in the amount of loaded drug has been confirmed, but with the most biomimetic formulation (Poly(AA-co-AM-co-NVP-co-HEMA-PEG200DMA)) a significant increase in loading is demonstrated yielding the greatest loading potential (the highest loading achieved to date and 6× over control networks due to multiple binding points with varying functionalities) (FIG. 5C).

II. Dynamic Drug Release Profiles

Dynamic release profiles in artificial lacrimal solution and an artificial lacrimal solution with protein, demonstrated extended release of a viable therapeutic concentration of ketotifen. Release studies confirmed that release rates can be tailored via type and amount of functionality and extended from one to five days. FIG. 6 highlights normalized data of the fraction of drug released versus time (mass delivered at time t divided by the mass delivered at infinite time). For poly(n-co-HEMA-co-PEG200DMA) networks (where n was AA-co-AM, AM, or AA), the release of drug showed a relatively constant rate of release for approximately 1 day, with little difference in the release profile. However, the most structurally biomimetic network, poly(AA-co-AM-co-NVP-co-HEMA-PEG200DMA), exhibited a five-fold increase in the extended release profile (i.e., approximately 5 days).

It is hypothesized that providing all the relevant functionality to the mimicked docking site with the proposed polymer synthesis technique affords a higher affinity of the drug for the network and thus an even slower release of drug compared to control networks. Furthermore, a five to seven day release profile fits quite well into the time usage of one-week extended-wear soft contacts.

It has been demonstrated that the loading of drug can be controlled by the type, number, and diversity of functionality within the network. The loading (and hence the mass delivered) can also be controlled by the initial loading concentration of the drug. We have demonstrated control over the cumulative mass of drug released by changing the loading concentration. By considering the relative size of our gels (i.e., gels were slightly bigger than normal lenses) and mass of drug released in comparison to typical ophthalmic eye drop dosages (ketotifen 0.25 mg/mL of solution with one drop every 8 hours), the preliminary results revealed that a therapeutically relevant dosage could be delivered for extended periods of time.

To investigate the effect of protein on dynamic release, we chose lysozyme as a model protein since it is the largest protein component in tear fluid. FIGS. 7A-B highlights the poly(AM-co-HEMA-co-PEG200DMA) network release profile in artificial lacrimal solution with lysozyme, which leads to a factor of 5 increase in the duration of release. For the most structurally biomimetic network, poly(AA-co-AM-co-NVP-co-HEMA-PEG200DMA), this could lead to a sustained release approaching 25 days. These studies demonstrate that the time of release may be delayed even further in an in vivo environment, leading to a substantial increase in applicability of contact lens ocular delivery.

III. Polymerization Reaction Analysis

The rate of polymerization for a given conversion decreased for increasing mole percentage of template molecule in pre-polymerization monomer solution. Thus, the formation of polymer chains and the enhanced loading due to the configurational biomimetic effect may be related to the propagation of polymer chains. The template molecule poses physical constraints to free radical and propagating chain motion and hence effectively lowers the rate of polymerization in the creation of ligand binding pockets. These results show that imprinting is reflected at the molecular level. For a given conversion, the rate of polymerization was lower for the multiple functional monomer pre-polymerization mixtures than the single monomer mixtures. We hypothesize that biomimetic imprinting with multiple monomers results in the formation of better ligand-binding pockets with enhanced loading properties which leads to slower rates of polymerization.

IV. Equilibrium Swelling Profiles and Mechanical Property Analysis:

Equilibrium swelling studies in DI water and 0.5 mg/ml concentrated ketotifen solution) indicated that recognitive and control networks were statistically the same and that 40% of the swollen gels is water, which indicates that the comfort of wearing and oxygen permeability of these gels is in agreement with conventional contact lenses. These studies indicated that biomimetic imprinting, and not an increased porosity or surface area of the gel, is responsible for the enhanced loading properties. It also demonstrated that the loading process does not affect the rate of swelling of the polymer matrix.

Further studies on the mechanical properties of the gels have shown comparable storage and loss moduli, glass transition temperatures and damping factors to that of conventional contact lenses (data not shown). Each gel produced was optically clear and had sufficient viscoelasticity to be molded in thin films (for refractive differences)

V. Synthesis of Fluconazole Recognitive Networks:

The functional monomers were selected by analyzing the binding mechanism of fluconazole with that of fungal cytochrome P450. The functional monomers chosen were n-vinyl pyrrolidone (NVP) hydroxyethyl methacrylate (HEMA) diethylaminoethyl methacrylate (DEAEM), and acrylamide (AM). The crosslinking monomer was polyethylene glycol dimethacrylate (PEG200DMA) and the initiator for the polymerization reaction was azobisisobutyronitrile, AIBN. Cross linker monomer, functional monomer, and templates were placed in an amber bottle and allowed to equilibrate in a light-controlled location. The initiator was added after the monomers and template were completely mixed. The solutions were mixed and used within 24 hours. The control and sample solutions were added to a template between two glass plates, and polymerized using UV polymerization in a nitrogen environment. The gels were washed and cut into uniform disks with a 10 mm cork borer. Binding studies were performed by placing washed gels into solutions of incremental drug concentrations and recording absorbance after equilibration to determine the amount of drug bound. Release studies were performed by placing the drug-laden gel into a lacrimal solution and taking absorbance readings at regular intervals to determine the release.

Fluconazole belongs to the azole class of antifungals and works by binding with the fungal cytochrome p45 lanosterol 14a-demethylase to inhibit fungal cell wall growth. By analyzing the Mycobacterium tuberculosis P450 CYP121-fluconazole complex, hydrogen bonding is provided by Threonine, Serine, and Glutamine residues with hydrophobic interactions from Valine, Phenylalanine, and Methionine.

FIG. 9 shows a graph 900 of a Fluconazole binding isotherm of 5% cross-linked poly-HEMA-co-AM-co-DEAEM-co-NVP-co-PEG200DMA hydrogel lenses. The recognitive hydrogel network is represented by the line 911 and the control hydrogel network is represented by the line 913. FIG. 10 shows a graph 1000 of Fluconazole dynamic release represented by the line 1001 in artificial lacrimal solution from 5% cross-linked poly-HEMA-co-AM-co-DEAEM-co-NVP-co-PEG200DMA hydrogel lenses.

VI. Forming a Living Polymer for Enhanced Drug Loading and Delay Drug Transport:

FIG. 11A illustrates a graph 1100 representing drug loading differences between free radial polymerization and living polymerization and iniferter reaction strategies. The control polymer network is represented by the line 1113, the recognitive polymer network is represented by the line 1111 and the living polymer network is represented by the line 1115. A typical polymerization solution resulting in polymer designated as Recognitive Gel (FIG. 11A), was made with 0.187 mL ethylene glycol dimethacrylate or EGDMA (0.993 mmole), 0.16 mL methacrylic acid or MAA (18.86 mmoles), 18.55 mg of azo-bis(isobutyronitrile) or AIBN, and 84.06 mg of EA9A template (ethyl adenine-9-acetate (EA9A)). Solutions were placed in a sonicator for several minutes until all solids were dissolved. The Control Gel solution was made with exactly the same formulation except no EA9A template was added. The polymer Recognitive Gel (Iniferter) was made by addition of iniferter, 3.89 mg TED. The molar ratio of the initiator to the iniferter in Recognitive Gel (Iniferter) was 8.61 with 0.187 mL ethylene glycol dimethacrylate or EGDMA (0.993 mmole), 0.16 mL methacrylic acid or MAA (18.86 mmoles), 18.55 mg of azo-bis(isobutyronitrile) or AIBN, and 84.06 mg of EA9A template. In these gels, the molar ratio of the initiator to the iniferter gave double bond conversions similar to the double bond conversions for the Recognitive Gel. For polymerization, the temperature of polymerization was (14° C.+/−1° C. throughout exothermic reaction). The molar ratio of initiator to iniferter can range 0.0001 and upward.

FIG. 11B illustrates a graph of fractional release of diclofenac sodium template for poly(DEAEM-co-HEMA-co-PEG200DMA) gels. The recognitive gel prepared via “living/controlled” polymerization (A) had further extended release in water when compared with the recognitive gel ( ) and the control gel (x). The error bars represent the standard error with n=3.

CONCLUSION

Polymerization kinetics in the presence of the template reveal mechanisms of interaction as well as provide criteria with which other template-monomer systems can be chosen experimentally. The use of a biomimetic approach for synthesizing recognitive hydrogel polymers has led to the development of an ophthalmic drug delivery system using contact lenses formed from the recognitive hydrogel polymer. The ophthalmic drug delivery system of the present invention can provide improved bioavailability and efficacy of drug delivery and exhibit controlled time release of the drug. The ophthalmic drug delivery system can be tailored to exhibit properties suitable for the intended drug therapy and has a potential to replace traditional eye drop therapies and other methods.

EXAMPLE 2 Hyaluronic Acid Introduction

Within the human body, Hyaluronic Acid (HA) binds to various receptors, the most significant is the cell-surface glycoprotein CD44. Using molecular modeling and site-specific mutagenesis, the amino acid residues most responsible for the binding of CD44 to HA Residues deemed critical for HA binding have been identified as Tyr-42, Arg-78 and Tyr-79. Residues considered important for HA binding were Lys-38, Arg-41, Lys-68, Asn-100, Asn-101 and Tyr-105.

Based on this analysis acrylate monomers were sought with functional groups that bore similarities to tyrosine, arginine, lysine and asparagine. Tyrosine contains a 4-hydroxypheny I group which features aromatic behavior with some hydrogen bonding capability. Arginine and lysine have amine groups which bear positive charges when protonated. Asparagine possesses an amide group for additional hydrogen bonding. While these monomers were used, other monomers of the requisite functional chemistry and amino acids as discussed previously or other biomolecules such as nucleic acid aptamers and more are able to provide recognition/delayed release.

Acrylate monomers with similar chemical behavior are acrylamide (AM), N-vinyl pyrrolidone (NVP) and diethylamino ethyl methacrylate (DEAEMA or DEAEM). AM shares an amide group with asparagine. NVP, an aromatic lactam, is able to be seen as an analog to tyrosine for its aromaticity and hydrogen-binding capability. Finally, DEAEMA is a cationic acrylate and is similar to arginine and lysine because of its positive charge.

These monomers, if incorporated into a hydrogel network, would non-covalently interact with HA and increase the affinity of the molecule for the hydrogel, thereby giving an additional level of control over the release rate. The DEAEMA was expected to form an ionic bond with the carboxylic groups on the gluconic acid units, and the AM and NVP to form general hydrogen-bonds with various groups on both gluconic acid and acetylglucosamine. The increased ability to tailor the release kinetics enable an optimum designed formulation for a desired product. HA within the lens is also able to lead to less protein adsorption and increased water content.

Biomimetic Imprinting. Biomimetic Imprinting (BIP) is a technique for synthesizing hydrogels and conferring certain unique physical properties to it in the process. BIP is based on the premise that synthesizing a hydrogel by crosslinking its constituent monomers in the presence of template molecules are able to alter the microstructure of the gel. Consequently, it is able to change the way the hydrogel interacts with the environment and with molecules diffusing through it. To prepare a hydrogel by BIP, the constituent monomers are mixed with a selected molecule and allowed to reach equilibrium. The monomers complex with the template molecules to lower the free energy of solvation in the mixture. If the hydrogel is cross-linked with the monomers in these optimum configurations, then the hydrogel retains a molecular “memory” of the template molecule which persists even if the template is washed away. This molecular memory enhances the affinity of the hydrogel for the drug used as the template.

The selection of monomers involved in the BIP process is critical for effective imprinting. This involves identifying binding molecules and other biological agents that have evolved to bind with the template molecule. The active site of the binding molecule is examined for amino acids that are critical for binding to the template molecule. Analogous acrylate monomers are selected by comparing the chemistry of the acrylates to the critical amino acids. The acrylate monomers with functional groups most similar to the critical amino acids are likely to form the best memory sites within the hydrogel. This technique has been demonstrated for the molecule ketotifen fumarate.

When the functional monomers are added to the HA-polymer backbone mixture and allowed to equilibrate with HA, the monomers would prefer to be spatially arranged in a low energy configuration. Such a configuration would favor electrostatic and polar interactions between the monomers and the HA, much like the interactions between amino acids and HA in the CD44 binding site. When the gel is cross-linked, the monomers would be immobilized in these favorable configurations, cresting sites within the network with a stronger affinity for HA than areas with the same chemical composition and random configuration.

Polymer Gel Synthesis. In BIP studies, acrylate monomers are incorporated into the hydrogel, with analogous chemistry to the described amino acids. The objective is to demonstrate control over the release kinetics of the HA from the hydrogel by varying the degree of affinity between the HA and hydrogel.

We added appropriate amounts of the functional monomers acrylamide (AM), N-vinyl pyrrolidone (NVP) and diethylamino ethyl methacrylate (DEAEMA). The ratio between AM, NVP, DEAEMA was kept at 1:1:2 for experiments matching the proportions of the corresponding amino acids found in the CD44 binding site. However, differing compositions and other chemistry with similar functionality are able to be favorable. Each formulation is identified by the percentage of the total prepolymerization mixture that is composed of added functional monomers. For instance, a formulation containing 99% of other monomers and macromers and 1% of added monomers (including AM, NVP and DEAEMA) is referred to as 1% monomer composition. Crosslinking of the hydrogel is done for 45 seconds with an intensity of 10.5 mW/cm².

Each hydrogel lens is synthesized by the following protocol: 5 g of other bemonomers/macromers (PM) (these constituents make up the backbone and crosslinking structure of the network are pipetted into a 15 mL centrifuge tube. HA in powder form is added in a concentration of 6.5 mg per gram with other macromers (PM). (Other concentrations of HA are able to be used as noted). The mixture is vigorously stirred and then centrifuged to remove air bubbles. The mixing and centrifugation processes are repeated until the HA is homogeneously incorporated into the prepolymerization mixture.

Molds for the lenses are contrived with two microscope slides (2″ by 1.5″) separated by a 125-micron thick Teflon spacer. The spacer matches the dimensions of the glass slides and has a central 1-inch square cavity. For stability, the spacer is taped to one of the slides.

A mass of 125 to 200 mg of the HA-incorporated mixture is pipetted into the central cavity of the mold. The second slide is placed on top to sandwich the mixture, and the two slides are clamped together with binder clips. The mold is exposed to UV light at an intensity of 10.5 mW/cm² for a duration of 15-45 seconds. Other initiating mechanisms or polymerization schemes are also able to be used such as thermal, controlled or living polymerization, radox, condensation, polymerization, and more. The slides are separated to expose a film of cross-linked hydrogel. The film is softened by soaking in 1-5 mL of DI water, lifted off the mold and cut into a disk with a cork borer (14 mm diameter). The disk is the therapeutic contact lens for the purposes of in vitro release experiments.

To prepare strips for the tensile studies, the Teflon spacer is cut with an inner space of dimensions (6 cm by 3 cm). Crosslinking takes place at an intensity of 10 to 12 mW/cm². The hydrogel is cut with a clean blade to strips 6 to 10 mm wide.

Dynamic Release Studies. Dynamic release studies were conducted to measure how long a therapeutic contact lens would release drug in vitro under conditions similar to the eye surface. The studies are also able to reveal the order of the release.

The protocol for a kinetic release study begins with the synthesis of hydrogel as described previously. As each study was conducted in triplicate, three lenses were made of every HA-prepolymer formulation. A lacrimal solution isotonic with human tears (6.78 g/L NaCl, 2.18 g/L NaHCO₃, 1.38 g/L Kcl, 0.084 g/L CaCl₂ 2H₂O, pH 8) was pipetted into 50 mL centrifugation tubes, 20 mL per tube. Each lens was placed in a tube and incubated at 35EC on an orbital shaker. After measured time intervals, the lenses were extracted and deposited into fresh lacrimal solution. The lacrimal solution with HA released in it was stored at 4° C. until essayed. After data had been collected for 3-5 days, an ELISA assay for HA was used to determine the concentration in each release sample. If necessary, the samples were diluted to bring into the assay detection range.

Cumulative release curves were plotted for each lens formulation, and the data was used to calculate the rate of diffusion of HA from the hydrogel. The transport of the drug from the lens placed in a fluid environment was mathematically modeled. As the length of the exposed surface of the lens was greater than ten times the thickness, the edge effects were ignored and the system was assumed planar with diffusion occurring in one dimension. Assumptions were made to apply the model.

Swelling Studies. Swelling studies were conducted on lenses with and without HA eluted. Lenses that had eluted HA were collected after the final release step and air dried for at least 24 hours. Lenses with HA were dried in a similar manner after synthesis. They were then transferred to a vacuum oven and dried at 30° C. and 28 inch Hg for a minimum of 5 days. Upon removal from the oven, the lenses were either stored in a desiccator or immediately weighed on a microscale balance, in air and also in n-heptane (a non-solvent to the hydrogel). The lenses were hydrated in DI water for at least 24 hours until fully swollen. They were again weighted in air and in n-heptane. By Archimedes' principle of displacement and with the specific gravity of n-heptane, the volume of the hydrogel lens in dry and swollen states was calculated. As controls, lenses with similar compositions but no HA underwent the same treatment.

From this state, a number of parameters were able to be calculated such as the equilibrium weight swelling ratio (q), equilibrium swelling ratio (Q) and polymer volume fraction (v₂). The equilibrium weight swelling ratio is the ratio of the weight of a fully swollen gel to the weight of a dehydrated gel. The equilibrium volume swelling ratio is the ratio of the volumes of the same. The polymer volume fraction is the ratio of the volume of the dry polymer to the same when fully hydrated.

Heat Stability Studies. Hydrogels were subjected to simulated sterilization conditions to determine if the release characteristics would be affected. PM or non-functionalized hydrogels (no functional monomers) were prepared with 6.5 mg HA/g PM, no added monomers, and placed in 2 mL microcentrifuge vials with 6.5 mg/mL solution of HA in DI water to prevent partitioning out of the HA. The pH of the solution was adjusted to 11. The vials were heated to 120° C. for 40 minutes and then cooled in a room temperature water bath. The lenses were removed from the vials, blotted to remove excess HA from the surface and studied for their release kinetics.

The effect of heat conditions on HA solution was also assessed. In 2 mL microcentrifuge vials, 1 mL samples of HA solutions of 500 ng/mL and 10 μg/mL were heated to 105° C. and 121EC respectively. The 500 ng/mL samples were heated for 0 minutes, 5 minutes, 30 minutes and 60 minutes while the 10 μg/mL samples were heated for 15 minutes, 30 minutes, 45 minutes and 60 minutes. The sample vials were quenched in a room temperature water bath and assayed with an ELISA assay kit for HA.

Studies were also completed to determine the distribution of HA fragments or chains after a sterilization time of 40 minutes at 120EC. A 6.5 mg/mL solution of HA in water and a 6.5 mg HA/g PM lens were subjected to the heating period and the chain fragments were analyzed via size exclusion chromatography using water mobile phase with PL aquagel-OH columns and standards of HA of varying molecular weights.

Dynamic Mechanical Analysis—Tensile Studies. Hydrogels prepared in strips (in triplicate) were mounted on a dynamic mechanical analyzer (RSA III, TA Instruments) at a gauge length of 30 to 35 mm, and extended at a rate of 4 mm/min. The gels were fully hydrated through the experiment, and hydration was maintained with an aerosol diffuser.

Optical Transmission Studies. Optical transmission studies were conducted by cutting small diameter films and placing in the bottom of a 96 well plate where absorbance values were measured via spectrophotometric monitoring. All films were fully hydrated and tested at wavelengths of visible light (380 to 780 nm). The absorbance value of each well in air was calculated and subtracted from the data. Percent transmission values were calculated from the absorbance data.

Molecular Size Effect on Release Rates. The fact that HA is a long chain molecule raises interesting questions about its manner of release. A small molecule would encounter relatively little hindrance from the polymer network as it diffuses out of the hydrogel. The diffusion coefficient would therefore be relatively high. As the molecular weight of the drug increases and the effective radius becomes comparable to the polymer mesh size, it faces more steric hindrance and the diffusion coefficient decreases.

If the mesh size of the hydrogel is much larger than the effective radius of the HA, then there is little ability to tailor the delivery of HA by modifying the hydrogel. To determine if the mesh size allows potential development of the technology, release studies were conducted using HA of different molecular weights. HA-PM lenses were synthesized with 6.5 mg of HA per g of monomer/macromer (PM), using HA of three different size: 1 million daltons, 100 kilodaltons and 50 kilodaltons. Crosslinking was done for 8.2 seconds with UV light at an intensity of 9.2 mW/cm².

FIG. 12 is a plot of the release profiles of HA of different sizes normalized to the maximum amount of HA released. It is evident that the smaller HA molecules diffused out of the hydrogel lenses rapidly, compared to the largest HA molecule which was diffusing out of the lens for up to 3 days.

The diffusion coefficients decrease with increasing HA size, indicating that the mesh size is comparable to the effective radius of the HA. Modifying the pre-polymers with appropriate chemical functionalities, additional control is able to be exerted over the release characteristics of the system. In the studies, unless noted, 1 million Dalton HA was used.

Concentration of HA in Hydrogel. The first objective was to ensure that the lens is capable of delivering a therapeutic dosage of HA over an acceptably long time period. Too low a dosage would be insufficient for treatment, while an excessive dosage would feel uncomfortable and possibly limit vision.

The amount of HA in a therapeutic dosage was determined by referring to commercial eye drop formulations of HA for artificial tears, In particular, Aquify® Long-Lasting Comfort Drops were examined, a 0.1% solution of HA. According to the package insert, the recommended dosage is 2 drops up to 3-4 times daily. The volume of a typical eye drop is 20 μL so if drops are delivered four times a day, the delivered dosage would be 40 micrograms of HA every 6 hours, assuming 100% bioavailability.

There is no easy way to predict a priori what the release rate will be for a given concentration of HA incorporated into the pre-polymer mixture. Three concentrations of HA-PM were selected, lenses were made and release studies were conducted to help benchmark the appropriate amount of HA to add to the formulation.

To synthesize the lenses, the protocol outlined previously was followed. At the step that involves dissolving HA into the monomers/macromers of PM, the protocol was deviated from in the following manner: in addition to preparing a sample by adding 6.5 mg per gram of monomers/macromers (PM), a sample was prepared of 2 mg per gram of monomers/macromers and a sample of 40 mg per gram of monomers/macromers (PM). The monomers/macromer-HA in the molds was cross-linked at an intensity of 9.2-10.5 mW/cm² for a duration of 8.2-15 seconds per sample.

FIG. 13 shows the cumulative HA releases from lenses made with formulations containing 2 mg/g, 6.5 mg/g and 40 mg/g of HA in monomers/macromers (PM). Release was measured over 24 hours. The 2 mg/g formulation released approximately 15 micrograms over the first 6 hours. The 6.5 mg/g formulation released 41 micrograms over the first six hours, and then decreased its release rate. The 40 mg/g formulation released a large amount of HA (approx. 581 micrograms for the first 6 hours. All curves exhibited a Fickian release profile. The 6.5 mg/g and 40 mg/g lenses released HA for at least 24 hours. The data has been analyzed in determination of the diffusion coefficient and it appears, as expected, that the diffusion coefficient is independent of the concentration of HA in the lens, and the release amount of HA depends on the concentration incorporated into the hydrogel. The 6.5 mg/g formulation gave the closest match to the desired HA delivery rate, thus was selected as the basis for future experiments. FIG. 14 highlights the release from only the 6.5 mg/g formulation.

Results:

I. Effect of Sterilization Conditions on Release of HA

Since the contact lenses are being developed for commercial use, it is important that there be a protocol for sterilizing the lens before application to the eye.

This method was reproduced in lab by preparing lenses with the formulation of 6.5 mg/g of PM (no added functional monomers). They were placed in a known volume of an aqueous solution of 6.5 mg/mL hyaluronic acid (pH 11) to minimize HA partitioning out of the hydrogel during the experiment. The lens and accompanying fluid were sealed inside a 2 mL microcentrifuge tube and heated in a sand bath at 120EC for 40 minutes. After cooling, the lenses were extracted, blotted to remove excess HA from the surface, and used to conduct release studies. From FIG. 15, it is clear that the release pattern of the basic formulation (with no added monomers) before and after heat sterilization is equivalent. It is possible that the high concentration of the HA in the lens as well as the polymer has the effect of protecting the HA from degradation.

II. BIP and Imprinted Early Studies—Effect of pH on Release of HA

Imprinted gels were prepared by matching HA biological binding information. The ratio between AM, NVP, DEAEMA was kept at 1:1:2 for experiments matching the proportions of the corresponding amino acids found in the CD44 binding site. FIG. 16 highlights early studies including the functional monomers in the formulation at a 1% functional monomer. The imprinted gels released very negligible amounts of HA. The possibility that the non-existent release of HA from the monomer-added PM hydrogel was because the HA molecules were being immobilized by the monomers reacting to form sterically immobilizing structures around them. If this were the case, the significantly decreased diffusion coefficient in the presence of functional monomers would not be because of electrostatic or hydrogel-bonding interactions with the HA. To examine this possibility, a series of release studies at various pH were conducted.

At neutral pH, the DEAEMA is protonated and positively charged while gluconic acid (a monomer of HA) is deprotonated and negatively charged. If the HA is non-covalently bonded to the hydrogel via electrostatic interactions between DEAEMA and gluconic aciden, then increasing the pH of the lacrimal solution would deprotonate the DEAEMA. The electrostatic interaction with HA would decrease and allow the HA to be released from the gel. If changing the pH does not increase the HA release rate, then it is unlikely that interactions between the DEAEMA and gluconic acid units are responsible for the delayed release of HA. FIG. 16 shows release data conducted on lenses made with the 1% formulation HA. At pH 8, the release of HA from the lenses was negligible. Similar results were obtained at pH 11. However, when pH was increased to 12, the HA release rate dramatically increased and the total amount release was slightly higher than the similar non-imprinted gel. At pH of 12 DEAEMA is not positively charged and the release of HA is seen in a 24 hour period. At pH 3 and 8, the charges exist on both HA and DEAEMA of varying degrees. There is significant potential to tailor release rate by altering monomer composition as well as charges.

This evidence lends weight to the theory that electrostatic interactions contribute the delayed release of HA from the monomer-added PM gels as neutral pH. An optimization of these interactions pre- and post-polymerization will be a tool to delay release.

FIG. 17 shows the release profiles of various formulations containing different weight percentages of functional monomers in the hydrogel. Once again, the ratio between AM, NVP, DEAEMA was kept at 1:1:2 for experiments matching the proportions of the corresponding amino acids found in the CD44 binding site. The release profile of HA-PM without added functional monomers is also shown as a control. The total amount of HA released depends on the composition of the prepolymerization mixture: formulations with higher percentage monomer compositions released less amounts of HA. A small change in the formulation mixture based on CBIP is able to tune the release profile.

The diffusion coefficients and orders of release for the different lenses were calculated (FIG. 18). There is a clear trend in the diffusion coefficients of the different formulation: as the percentage of functional monomers increases, the diffusion coefficient of HA decreases (FIG. 19). FIG. 19 also demonstrates the diffusion coefficient dependence on the polymer volume fractions, which have been shown to correlate to gel structural parameters. To eliminate the possibility that the variation in diffusion rate stems from a change in overall structural parameters, the diffusion rates were plotted versus the polymer volume fraction. From the graph it is clear that the polymer volume ratio remains fairly constant (0.244+/−0.003) while the diffusion coefficients vary dramatically.

A strong linear relationship becomes apparent when the diffusion coefficients are plotted versus the monomer percentage (FIG. 20). Extending the trendline indicates that when the percentage composition is at or above 0.33%, the diffusion coefficient approaches zero. This agrees with the release data for the 1% and 5% monomer composition lenses; both released negligible amounts of HA.

III. Diversity of Functional Groups

Increasing the variety of monomers incorporated into the hydrogel is able to produce memory sites within the structure that are likely to bear a chemical resemblance to the binding sites in CD44. According to this hypothesis, adding AM, NVP and DEAEMA to the formulation together would lead to a higher affinity for HA than in the case of DEAEMA alone or AM-NVP together without DEAEMA. A set of lenses with varying proportions of the functional monomers AM, NVP and DEAEMA, with the total amount of functional monomer fixed at 0.125% of the total prepolymerization mixture. Following the format [AM:NVP:DEAEMA], the proportions used in the formulations were [1:1:0] (equal parts AM and NVP), [1:1:1] (equal parts AM, NVP and DEAEMA), [1:1:2] (equal parts AM and NVP, and two parts DEAEMA—this is the proportion used in all of the earlier presented formulations, and [0:0:1] (only DEAEMA). The hydrogels were cross-linked for 45 seconds with an intensity of 10.5 mW/cm². Release was conducted by following the replacement protocol.

FIG. 21 highlights two formulations with 0.125% DEAEMA. One has only DEAEMA and the other has AM and NVP included at equal proportions of 0.0625%. The results suggest that the addition of AM and NVP is significant and decrease the transport of HA from the gel.

The fractional release of all diversified gels is shown in FIG. 22. The results indicate that an overall functional monomer percentage greater than 0.125% is needed to alter the shape of the release curve. It may be highly possible to push this towards zero order release moving closer to 0.33% functional monomer percentage composition.

FIG. 23 shows the influence of the different functional monomers on the HA diffusion coefficient. It is clear that multiplicity of interactions significantly decreases the diffusion coefficient of HA from the lens. Also, increasing the number of these monomers further decreases the diffusion coefficient. At the extremes of the curves, where either AM and NVP reside or only DEAEMA, it is clear that each exerts some effect on the delayed transport.

IV. Swelling Studies

The possibility exists that the addition of monomers to the formulation is able to change the polymeric structure of the resulting hydrogel. If the structure changed such that the size of the mesh between polymeric strands became smaller, the HA would encounter more barriers to diffusion and have a lower release rate. Therefore, it is needed to distinguish whether the decrease in release rate was caused by the chemical functionality of the added monomers or by a decrease in mesh size.

FIGS. 24 and 25 show how the polymer volume fraction varies with respect to various parameters, including the percentage of monomers added to the formulation both in the presence and the absence on HA, the sterilization process, and the percentage of added functional monomers. The polymer volumes fractions fail within a narrow range of values, 0.24 to 0.3, which are very close to the specifications of the pure polymer (no imprinting) (0.31).

FIGS. 26 and 27 show that the equilibrium weight and swelling ratios and polymer volume fraction in the swollen state of gels pre-HA release were equivalent to gels that were studied after HA release. Studies of most gels (e.g. gels without an indication of ‘unreleased’) were conducted after the incorporated HA was released from the network.

V. Structural Analysis

There are many theoretical models that enable obtaining structural and configurational information about a hydrogel from experimental data obtained by swelling and tensile studies. In particular, information is able to be obtained about the mesh size of the hydrogel and determine whether a tighter mesh is responsible for the decrease in the diffusion coefficient of HA through imprinted hydrogels.

The Peppas-Merrill model describes the relationship between the average molecular weight between crosslinks and the equilibrium polymer volume fraction (v₂) for a swollen hydrogel cross-linked in the presence of a solvent.

Equilibrium swelling studies were performed on the hydrogels to determine the weight swelling and volume swelling ratios (q and Q) and the polymer volume fraction (v₂). The equilibrium swelling ratio, Q, is the ratio of swollen to dry volumes of the hydrogel. The weight swelling ratio, q, is the ratio of swollen to dry weights. The polymer volume fraction in the swollen state, v₂ is the ratio of the dry polymer volume to swollen polymer volume, and (1−v₂) gives the fractional water content of the hydrogel. These parameters were calculated for all hydrogels with some values presented in FIG. 25. The hydrogels synthesized with HA generally have slightly higher Q and q than the hydrogels synthesized without HA. The former also have lower v₂, indicating higher water content. Two factors influence this difference. First of all, the presence of HA in the prepolymer mixture is able to influence the formation of polymer chains and associated crosslinking points, making the polymer chains more mobile and increasing the hydrogels' capacity to hold water. Secondly, the residual HA in the hydrogels increases the hydrogels' capacity to hold water. HA within the lens is also able to lead to less protein adsorption and increased water content.

In general, all of the hydrogels have v₂ falling within the range of 0.23 and 0.29 suggesting that the mesh size is similar for all of the gels synthesized. In particular, it is noted that the swelling parameters of hydrogel with 1%-by-mass functional monomers remain the same at pH 7 and pH 12, indicating that the pH dependent increase in HA release did not result in change in mesh size. Furthermore, the swelling parameters do not charge for gels synthesized with HA despite heat sterilization.

In FIG. 19, the relationship between the diffusion coefficient and polymer volume fraction in gels with various %-by-mass of functional monomers is illustrated.

The figure clearly reveals the narrow range of polymer volume fraction of the hydrogels. In contrast, the diffusion constants of HA through these networks varies dramatically. The highest diffusion coefficient (from the gel without functional monomers) is nearly 4 times higher than the lowest diffusion coefficient (from the gel with 0.25% functional monomers). This is strong evidence that the diffusion coefficients do not vary because of structural parameters such as the mesh size.

To determine mesh size, further structural analysis was done through tensile testing of the sample, in which the hydrogel samples were extended at a constant rate and the tension on the sample was recorded. The extension versus applied force were plotted for four types of hydrogel samples: gels without added functional monomers or HA (PM), PM with HA, PM with 0.25% by mass functional monomers, and PM with HA and 0.25% by mass functional monomers. Mechanical parameters for the hydrogels were obtained and summarized in Table 1 (gel mechanical parameters).

TABLE 1 Young's Shear Modulus Std. Hydrogel Modulus (M Pa) Std. Dev. (M Pa) Dev. PM 0.557 0.023 0.201 0.011 PM with HA 0.423 0.060 0.153 0.022 PM with 0.25% 0.535 0.032 0.195 0.015 f.m. PM with 0.25% 0.550 0.015 0.203 0.009 f.m. and HA

Of the four gels, three have very similar structural parameters. PM with HA (no functional monomers) has a modulus that is lower than the other hydrogels, indicating that the presence of the functional monomers decreases the mesh sizes slightly, but not to an extent that it is outside comparison with PM and PM with now added HA or functional monomers.

VI. Optical Analysis

FIG. 28 highlights the optical transmission studies of swollen hydrogels. The data indicates that the percent transmission of visible light through the gels is comparable. Absorbance and percent transmittance values (in triplicate) were uniform across the range of wavelengths studied (380 to 780 nm).

The percent change in concentration over time was lesser for the more concentrated solution (10 μg/mL) than the more dilute solution (500 ng/mL). The change in concentration in the 10 μg/mL solution was 30% over 60 minutes while the change in signal strength in the 500 ng/mL solution was 50% over 60 minutes. The percentage change in concentration versus time of heating is shown in FIG. 29. When solutions of HA in water are heated to above 100EC, they undergo some heat degradation. In the less concentrated solution (500 ng/mL) nearly 50% of the large HA degrades to shorter HA over 60 minutes, whereas in the higher concentration solution (10 μg/mL), the degradation is only 30% in the same time. This suggests that higher concentrations have a protective effect on the large HA molecule.

Thus, the higher concentrations of HA are able to have protective effects on the stability of the long-chain HA molecule. A dynamic release study assessed the heat effects on the HA hydrogel lenses. Comparing the heat-treated lenses with the control untreated lenses, similar release profiles were measured, suggesting that heat-treatment does not denature the HA within the hydrogel. The release profiles are shown in FIG. 30.

FIG. 31 illustrates a flowchart of a method of making a comfort agent delivery system. In the step 3100 a recognitive polymeric hydrogel is formed. In the step 3102, the recognitive polymeric hydrogel is formed into a contact lens with a comfort agent. As described above with reference to FIG. 1, additional steps are able to be included.

FIG. 32 illustrates a flowchart of a method of dispensing a comfort agent in an eye biological tissue. In the step 3200, a recognitive polymeric hydrogel matrix impregnated with the comfort agent is formed. In the step 3202, the recognitive polymeric hydrogel is placed in contact with a biological tissue to dispense the comfort agent. As described above with reference to FIG. 1, additional steps are able to be included.

The delivery of a comfort agent such as HA in a sustained manner is able to apply to, but is not limited to, any drug delivery device or coating where one is limited on space or volume of delivery device. Delivery of the comfort agent is able to apply, but is not limited to, particles, cylinders, spheres, films and coatings. Areas where delivery of the comfort agent is able to be applied includes, but is not limited to: refractive and non-refractive (cosmetic) contact lenses and ocular implant devices, bandage contact lenses—refractive and non-refractive (cosmetic) contact lenses—that are placed after major or minor eye surgery to release HA during the healing process and/or aid in ocular health, implant coatings, transdermal and drug delivery films, tissue engineering and regeneration materials (including in vitro) bandages, wound healing and tissue regeneration materials.

CONCLUSION

Polymerization kinetics in the presence of the template reveal mechanisms of interaction as well as provide criteria with which other template-monomer systems can be chosen experimentally. The use of a biomimetic approach for synthesizing recognitive hydrogel polymers has led to the development of an ophthalmic drug delivery system using contact lenses formed from the recognitive hydrogel polymer. The ophthalmic drug delivery system of the present invention can provide improved bioavailability and efficacy of drug delivery and exhibits controlled time release of the drug. The ophthalmic drug delivery system can be tailored to exhibit properties suitable for the intended drug therapy and has a potential to replace traditional eye drop therapies and other methods.

This technology creates a new architecture in polymeric films to enable enhanced drug loading and delayed time release of drugs. We have applied this to contact lenses to produce a new drug delivery system comprising contact lenses. These polymeric films are based on the creation of polymeric networks from a fundamental analysis of the biological recognition or biological mechanism of action of the therapeutic. The technologies produced from this work involve thin films for drug delivery, which can be applied to other areas besides the ocular market. This can apply to, but is not limited to any drug delivery device or coating where one is limited on space or volume of delivery device. It can apply, but is not limited to particles, cylinders, spheres, films and coatings.

Areas where this technology is applicable include, but are not limited to: Refractive and Non-refractive (Cosmetic) Contact Lenses and Ocular Implant Devices; Bandage Contact Lenses—Refractive and Non-refractive (Cosmetic) Contact Lenses—that are placed after major or minor eye surgery to release medication during the healing process and/or aid in ocular health; intra-ocular lenses; Nasal Drug Delivery via Ocular Lenses or Devices; Implant Coatings (Orthopedic, Adhesion prevention, etc.); Cardiovascular Device and Stent Coatings with Drug Release; Transdermal Patches and Drug Delivery Films; Tissue Engineering and Regeneration Materials (including in vitro); Bandages, Wound Healing;

Tissue Regeneration Materials; and Drug Tablet Coatings.

The technology also provides an opportunity for a combination drug platform where multiple drugs are released from a single lens of film. In some cases, multiple drugs of a certain type (combination anti-fungal therapy) and multiple classes of drugs (anti-inflammatory and antibiotic) are released to achieve more convenient and efficacious therapy.

The systems of the present invention lead to delayed transport in non-swellable (e.g., conventional contact lenses) but they may affect swellable hydrogel networks in a similar manner. It is believed that in the system of the present invention that a multiplicity of the interactions (multiple functional monomers) is preferred for optimal sustained release of a drug.

EXAMPLE 3

Comfort Molecule Delivery: Contact lenses have been shown to affect the natural flow profile of the tear layer due to a difference in surface tension. This disruption can cause a decrease in the rate of lacrimation and increase the rate of tear fluid evaporation causing a condition known as contact lens induced dry eye (CLIDE). Several recent surveys have shown that up to 80% of contact lens wearers report discomfort caused dryness. Of those surveyed, 30% report severe discomfort and more than half expressed dissatisfaction with their current lenses. Lens discomfort caused by CLIDE has been reported as a leading cause for the discontinuation of lens wear.

Administration of macromolecular re-wetting or comfort agents via eye drops is currently the most used method to overcome CLIDE discomfort. Re-wetting agents absorb water from the tear film and during blinking water is forced back onto the eye anterior surface, replenishing the tear film. In addition, the viscosity of the tear fluid can be altered by the presence of the macromolecule reducing the shear stress caused by blinking. However, drop volume, delivered dose, and patient compliance are highly variable with eye drop administration, and eye drops typically need to be instilled several times a day to maintain effective concentrations. For re-wetting eye drops, patients are advised to use drops as needed, and this can lead to extensive, inconvenient use with significant variations in comfort. Therapeutic molecules delivered via eye drops suffer from very low residence times with poor bioavailability and result in pulsatile concentration profiles in the tear fluid. Therapeutic soft contact lenses have been demonstrated to be an ideal platform for the controlled delivery of numerous drugs as well as comfort molecules. Producing more comfortable contact lenses or comfort-enhanced lenses is a significant goal, and this can be achieved by sequestering comfort molecules in the lens structure and/or releasing comfort molecule from the lens. Sequestering and/or releasing comfort agent can lead to less protein and lipid deposition and binding within the lens and increased lens water content, to less protein and lipid deposition on the surface of the lens, and to comfort agent in the tear fluid limiting evaporation and/or increasing tear water content.

Soft contact lenses are the most convenient platforms to alleviate CLIDE symptoms, especially when compared to alternative methods of treatment such as eye drops, inserts, and ointments. Incorporation of re-wetting agents into commercial lenses has been shown to make lenses more comfortable. This has been done by physically sequestering re-wetting agents within the lenses. Commercial lens product lines, such as Johnson and Johnson's Acuvue lens series and Coopervision's Proclear lenses, have incorporated polyvinyl pyrrolidone (PVP) (MW between 50-100 KDa) into the lens that does not release. The PVP incorporated in the lens lowers the dehydration rate of the lens. Cross-linked hyaluronic acid (HA) networks within lenses can also reduce dehydration and protein adherence to the lens providing dual mechanisms of comfort. Phospholipids can also be covalently attached to lenses to increase the ocular bio-tolerance and decrease the disruption of the tear film.

While sequestering re-wetting agents inside the lens to increase water content of the lenses can promote comfort, controlled release of the re-wetting agent may offer several advantages. The release of the re-wetting agent can provide a comfort effect on the surface of the eye away from the lens, refresh the lens surface, and increase the effectiveness of the re-wetting agent. Released re-wetting agent mimics delivery from eye drops, but it can occur at a continuous rate, providing similar comfort at lower concentrations taking the dose out of the patient's hands for optimal efficacy and convenience. It is probable that controlled release of re-wetting agent can provide greater comfort in dry eyes than lenses with sequestered comfort agent. The invention disclosed herein uses molecular imprinting for the controlled release of HA and/or other comfort agents from hydrogel contact lenses, beneficially in continuous extended wear lenses.

Development of silicone hydrogels allowed the production of continuous, extended wear lenses. Silicone hydrogel lens materials have been approved by the FDA for long durations of wear. For example, Air Optix® NIGHT & Day® Aqua silicone hydrogel lenses may be worn for up to 30 nights of continuous wear. Extended comfort in these types of lenses is extremely important since they reside on the surface of the eye for long periods of time, and extended comfort molecule release may be the best option for these types of lenses. To achieve 30 day release (or greater), application of controlled release methods through the molecular imprinting composition according to the invention are needed.

Molecular imprinting applied to silicone hydrogel contact lenses can significantly delay release and enhance comfort molecule loading. The invention discloses the first extended wear silicone hydrogel contact lens with extended, controllable release of hydroxypropyl methylcellulose (HPMC) using molecular imprinting. The controlled and extended release of 120 KDa HPMC from commercial silicone hydrogel lenses is demonstrated, wherein HPMC is a common re-wetting agent and rheology modifier used in over the counter eye drops.

Notably, there has been very little work highlighting the release of comfort molecules from silicone hydrogel lenses. Previous work involving silicone hydrogel lenses has not demonstrated significant loading or extended release of comfort agents. Soaking a variety of both silicone hydrogels and hydrogel commercial contact lenses in a HA solution did not load sufficient amounts of HA and release was complete in well under an hour. Work with much lower molecular weight molecules showed release was generally quick with the majority (i.e. 60-80%) of the loaded molecules eluting within 1 hour [Karlgard & Moresolt, Eye and Contact Lens. 2003; 29(2):83-89; Karlgard et al., International Journal of Pharmaceutics. 2003; 257:141-151] or within 3-4 days [Peng & Chauhan, 2010; 31:4032-4047; Kim et al., Journal of Controlled Release. 2010; 148(1):110-116; Kim et al., Biomaterials. 2008; 29:2259-2269].

Materials and Methods: Lotrafilcon B is a patented contact lens material used in CIBA Vision, Inc. lens brands such as Air Optix and others. The material is a mixture of the proprietary silicone hydrogel macromer, Betacon Macromer, referred to hereafter as macromer, methacryloxypropyl-tris-(trimethylsiloxy) silane (TRIS), and dimethyl acrylamide (DMA). The thickness of the commercial LFB lens varies depending on the power of the lens, but can be approximated as 100 μm center (swollen) thickness. The water content of the lens is 33%. The low water content and hydrophobic nature of the TRIS and macromer limits the maximum loading and potential re-wetting agents that can be loaded into the lens. An effective re-wetting agent must be hydrophilic to provide comfort and sequester water, but an agent displaying too high a preference for water will not disperse into the lens formulation and will not controllably release from the lens. For this reason, hydroxypropyl methylcellulose (HPMC) was selected as the most promising re-wetting agent as it demonstrates a wide range of solubility.

DMA was purchased from Sigma Aldrich (Milwaukee, Wis.), and TRIS and macromer were provided by CIBA Vision, Inc. All components were used as received. Darocur 1173, the photo-initiator, was purchased from Sigma Aldrich (Milwaukee, Wis.). Acrylic acid (AA), poly(ethylene glycol) (200) dimethacrylate (PEG200DMA), ethylene glycol dimethacrylate (EGDMA) and hydroxypropyl methylcellulose (HPMC) (MW=10, 90, and 120 KDa) were purchased from Sigma Aldrich (Milwaukee, Wis.) and used as received.

Synthesis of LFB Lenses and Synthesis of Imprinted LFB Lenses. Lenses were produced by mixing individual monomers of the LFB formulation. The monomers were kept refrigerated at 4° C. A typical formulation consisted of 1300 mg of Betacon macromer, 1300 mg of TRIS, and 1300 mg of DMA (referred to as the base formulation). The base formulation was equal parts of these monomers and used to calculate ratios of other components. Equal weights of EGDMA and PEG200DMA were added to a desired concentration anywhere between 0 and 10 wt % of the base formulation Imprinting monomers were then added between 0 and 10 wt % of the base formulation. Darocur 1173 was added to the solution to a concentration of 1% of the base formulation. The samples were thoroughly mixed by sonication for 30 minutes and then the comfort agent was added to the desired concentration. The formulations were exposed to high shear mixing for up to 1 minute and sonicated for at least 15 minutes to remove any dissolved air or air bubbles. After the comfort agent was added, the sample was sonicated and mixed until evenly dispersed.

A fixed volume was pipetted into polypropylene (PP) lens molds provided by CIBA Vision, Inc. The mass of formulation pipetted varied between 100 to 200 mg depending on the concentration of comfort agent. The lens was polymerized via UV polymerization using a UV light source (Novacure 2100, Exfo) with an intensity of approximately 25 mW/cm2 for a duration of 1.5 minutes. Out of mold lenses were −220 μm thick (center thickness) and water swollen lenses were −350 μm thick (center thickness) unless otherwise noted. Thinner lenses were produced with a swollen center thickness of ˜100 μm using a Thomas spherical joint pinch clamp tightened around the mold. The lenses were then removed from the mold and used immediately in optical clarity, swelling, and dynamic release studies.

Optical Clarity Studies. Optical clarity studies were conducted by measuring the percent transmittance of visible light (wavelength range from 450-750 nm) through swollen lenses. Lenses from the molds were cut with a No. 3 cork borer and placed in the bottom of a 96 well plate where absorbance values were measured via spectrophotometric monitoring (Biotek, Winooski, Vt.). Each lens was cut into at least three samples to note any local differences within the lens. All films were fully hydrated in 200 μL, of water with care taken that the film was in full contact with the bottom of the well plate and no air bubbles were present. The absorbance value of each well in water was calculated and subtracted from the data. Percent transmission values were calculated from the optical density and absorbance data. All optical clarity experiments were carried out in triplicate with three separate lenses, producing nine samples for each lens type.

Dynamic Mechanical Analysis—Tensile Studies. Hydrogels prepared in strips were mounted on a dynamic mechanical analyzer (RSA III, TA Instruments) at a gauge length of 30 to 35 mm, and extended at a rate of 4 mm/min. The hydrogels were molded to a sample specimen measuring 13 cm×13 cm×0.125 μm. The gels were fully hydrated through the experiment, and hydration was maintained an aerosol diffuser. The modulus was calculated from the initial slope of the stress vs. strain curve.

Swelling Studies. Equilibrium swelling studies were conducted by placing lenses in DI water until they reached equilibrium. The lenses were then removed from the solutions, patted dry with a soft tissue (Kimwipes®), and weighed in air using a balance (Sartorius). Dry weight values were calculated directly after removing from the mold. The equilibrium weight swelling ratio was calculated by the difference between the weight of the fully swollen lens and the weight of the dry lens divided by the weight of the dry lens.

Dynamic HPMC Release Studies. Dynamic release studies were conducted to measure how long each therapeutic contact lens would release HPMC in-vitro. The studies were conducted with the perfect sink model in order to sustain the greatest driving force, using a Sotax Dissolution Apparatus (Horsham, Pa.), in which loaded lenses were placed in 250 mL of DI water (pH 6.4). Separate experiments were conducted to confirm the perfect sink volume.

In the Sotax apparatus, the release media was stirred at a constant rate of 30 rpm by paddles and kept at a constant temperature of 34° C. The average weight of the lenses out of the mold was 35 mg±6 mg. HPMC concentration in the release medium was determined via HPLC (Shimadzu, Japan) equipped with a refractive index detector. The mobile phase was deionized water, and a flow rate of 1 mL/min was maintained by the HPLC. A standard curve of refractive index and known HPLC concentration was established. Percent mass release curves were plotted for each lens.

Microfluidic Device. A microfluidic device was created by mixing Sylgard 184 silicone elastomer base and curing agent in a 10:1 ratio. The mixture was stirred for 3.5 minutes and poured onto a glass plate within a circular mold. Within the mold were two needles to create apertures in the device for flow and one hemisphere of a glass marble (created from cutting an 18 mm diameter marble into a segment of 16.5 mm wide×5.7 mm high). The device was then cured at 60° C. under a vacuum for 6 hours. The PDMS was removed from the mold and then the needles and marble were removed. A slightly smaller marble hemisphere (15.2 mm wide×4.35 mm high) was placed under the mold and a syringe pump was used to pump DI water through the mold at the physiological flow rate of 3 μL/min. The inner chamber contained 175 μL of DI water. The lens was placed in the device, sealed, and the concentration was measured every 10-12 hrs. A schematic of the microfluidic device is presented in FIG. 33.

Results

FIG. 33 shows a Microfluidic Device Schematic. The device mimics the in-vivo drug release profiles by duplicating the physiological flow rate of the eye. The inner chamber is formulated by curing Sylgard 184 silicone elastomer around a glass marble. The specifications of the inner chamber are (A) 5.7 mm high and (B) 16.5 mm wide. The lens is lightly glued to a glass marble and placed under the mold. The measurements of the marble are (D) 15.2 mm wide and (C) 4.35 mm high. The inner chamber contains 175 μL of fluid. Two needles of (E) 0.125 μm diameter served as the inlet and outlet streams. The arrows represent the flow of DI water through the device with continuous flow of fluid over, around, and under the lens.

FIG. 34 shows Equilibrium Weight Swelling Ratio of LFB Lenses Containing 120 KDa HPMC. (A) Without Additional Crosslinkers (B) With Additional Crosslinkers. LFB lenses were synthesized in the presence of various concentrations of 120 KDa HPMC (pre-polymer loaded). The thickness of the lenses were (♦) 350 μm and (▪) 100 μm. N=3, and T=22° C. As HPMC content increases, water uptake in the lens increases. The LFB lenses prepared with additional cross linkers show significant decrease in water uptake, indicating therapeutic amounts may be loaded without excessive swelling.

FIG. 35 shows Dynamic Release of 120 kDa HPMC from LFB Lenses Without Imprinting. Dynamic HPMC release in DI water from 350 μm thick LFB lenses with various concentrations of pre-polymer loaded 120 KDa HPMC. Increasing the HPMC content in the lenses from (♦) 600 μg, (▪) 750 μg, to (▴) 3,000 μg HPMC/lens increased the overall release rate by increasing the concentration driving force and increased swelling. Optical clarity of the lenses were poor (i.e. less than 40% transmittance) due to significant swelling. N=3, T=34° C.

FIG. 36 shows Optical Clarity of LFB Lenses: (A) The Effects of Swelling and (B) Imprinting on LFB Lenses Prepared with Additional Crosslinkers. There is an inverse relationship between optical clarity and swelling. A comparison of 350 μm thick lenses prepared with various concentrations of pre-polymer loaded 120 KDa HPMC: (▴) LFB lenses, (▪) LFB lenses prepared with additional cross linkers and (♦) HPMC-imprinted LFB lenses prepared with additional cross linkers. Additional cross linkers increased optical clarity with imprinting also increasing clarity. Cross linker composition was a 50:50 wt. ratio of EGDMA and PEG200DMA. N=3, and T=22° C.

FIG. 37 shows Dynamic Release of HPMC from LFB Lenses: Effect of the Addition of Crosslinkers on HPMC Release Rate. Crosslinking monomers (xLer) were added to the LFB formulation to control lens swelling of HPMC-laden lenses. Lens compositions were: xLer/T ratio of 0 with 2,750 μg HPMC (♦), xLer/T ratio of 0.019 with 2,800 μg HPMC (▪), xLer/T ratio of 0.237 with 2,700 μg HPMC (▴), xLer/T ratio of 1.02 with 2600 μg HPMC (•), and xLer/T ratio of 1.75 with 2,600 μg HPMC(X). Crosslinks limited the free volume in the lens, thereby altering HPMC release and limiting lens swelling. The presence of cross linkers in the lens led to significant improvements in clarity with reduced swelling, but they are not sufficient to properly control HPMC release. By increasing the crosslinking monomer to template ratio (xLer/T), HPMC release is delayed by the decreased mesh size. As the percentage of cross linker is increased, there is a longer initial delay in release. However, there is not much control over the release rate after the initial delay. This is indicated by the equivalent rates of release once approximately 5-10% fractional release is achieved. Note: No functional monomers were present in these lenses. N=3, T=34° C.

FIG. 38 shows Extended Release of HPMC from Imprinted LFB Lenses: Effect of M/T Ratio on Release of 120 KDa HPMC. By using molecular imprinting and altering the M/T ratio, we can control the HPMC release rate. Increasing the M/T ratio decreases the release rate and can lead to a desired release rate. The rate of release for M/T=0, no additional cross linkers loaded with ˜3,000 μg HPMC/lens (▪) is <1 day to release ˜85%, M/T=0, loaded with ˜2,600 μg HPMC/lens (▴) is ˜260 μg/day for 10 days, M/T=0.2 loaded with ˜2900 μg HPMC (▴) is ˜300 μg/day for a duration of 13 days, M/T=2.8 loaded with ˜2200 μg HPMC (x) delivers ˜97 μg/day for 22.5 days, and M/T=3.4 loaded with ˜2900 μg HPMC(•) releases ˜60 μg/day for 53 days. All lenses had a xLer/T value of 1.5. Swollen lenses are ˜350 microns thick, T=34 C, N=3.

FIG. 39 shows Extended Release of HPMC from Imprinted LFB Lenses: Effect of M/T Ratio on Average Release Rate of 120 KDa HPMC. Release rate can be controlled by imprinting (M/T ratio). By extrapolating, there may be a region where HPMC is sequestered without release (molecular Velcro™ effect) due to a large number of functional monomer/HPMC interactions. All samples represented have Xler/T ratio between 1-2 and ˜2,500 μg HPMC. Swollen lenses are ˜350 microns thick, T=34° C., N=3.

FIG. 40 shows Optical Clarity of 100 μm Thick LFB Lenses Imprinted with 120 kDa HPMC. There is a wide range of compositions that will lead to an optically clear lens. Transmittance values are observed (▪)≧90%; (▪)≧80%; (▪)≧70%; (▪)≧60%; (▪)≧50%; (▪)≧40%; (▪)≧30%; (□)≧20%. The release and optical clarity data indicates that the cross linker amount (xLer/T) can be decreased.

FIG. 41 shows Dynamic Release in Physiological Flowrate of HPMC from 100 μm Thick LFB Lenses Imprinted with 120 kDa HPMC. Lenses contained 975 ug HPMC/lens with M/T ˜3.5 and xLer/T ˜1.5. The flow rate of water was 3 μL/min. The release was linear and constant with an average release rate of ˜16 μg HPMC/day. This represents the lower boundary of the release rate and can be increased to a desired rate. N=2, T=22° C.

I. Drug Loading and Performance

The first step of this work was to assess the feasibility of loading HPMC directly into a lens. This is the ideal case as it does not affect the synthesis or manufacturing process. Synthesizing an optically clear lens and then loading HPMC by soaking in an aqueous HPMC solution did not result in a sufficient reservoir of HPMC loaded into the lens to deliver comfort to the eye. The mass uptake of various HPMC solutions is presented in Table 2 (showing mass loading of various HPMC concentrations and molecular weights).

TABLE 2 Release Time Mass Loaded from Drug Solutions (μg/lens) After 7 Rewetting 1 Day Soak 2 Day Soak 7 Day Soak Day Agent 0.1% 1% 10% 0.1% 1% 10% 0.1% 1% 10% Soak HPMC 10 KDa 0.32 4.1 5.4 0.21 3.2 6.9 1.2 5.7 2.9 30 min 90 KDa 0.09 0.07 0.4 0.09 0.65 0 0.01 0.48 0 None Detected 120 KDa  0.001 0.003 0.004 0.08 0.04 0.01 0.006 0.2 0.1 None Detected

Loading of the 120 KDa HPMC did not exceed 0.2 μg/lens after soaking in a solution for a week. The highest loading was seen for 10 KDa HPMC. As the smallest molecular weight, the least resistance is seen to diffusion of HPMC in the lens. Release occurred in less than 30 minutes for the 10 KDa case, and for molecular weights above 10 KDa, very little HPMC release occurred.

The values are very low from a drug delivery perspective and indicate that soaked lenses are not a feasible option for the development of therapeutic contact lenses for long chain molecules. Loading of HPMC through adsorption is best explained as a thermodynamically unfavorable process. HPMC loaded into the hydrogel network would be more confined and adapt a much different conformation from the HPMC in the surrounding solution. The energy required for high molecular weight molecules to adapt a conformation change at the surface of the hydrogel and reptate deeply into the lens makes it unfeasible as a loading option. For smaller molecular weight molecules some loading may be demonstrated for hydrophobic molecules within silicone hydrogel lenses, but release will be short in duration without control leading to a Fickian or exponentially decaying release rate.

II. Optical Clarity and HPMC Dispersion

Dispersing the HPMC into the pre-polymer formulation, referred hereafter as pre-polymer loaded, and synthesizing the lens resulted in a sharp decrease in optical clarity of the lens, but a high concentration of HPMC could be dispersed into it. The maximum mass of 120 KDa HPMC loaded into a 350 μm thick lens was found to be ˜8,000 μg/lens. Beyond this concentration, lenses could not be formed, and LFB did not form the continuous phase in the lens. Qualitative mechanical property and optical clarity measurements led to the selection of 3,000 μg/lens as the optimum concentration of 120 KDa HPMC for 350 μm thick lenses. However, the optical clarity dropped from 92% transmittance in the unloaded swollen lens to 10% in the loaded swollen lens. The equilibrium weight swelling ratio approached values of 15 as HPMC concentration increased to 6,000 μg HPMC/lens in 350 μm thick lenses (FIG. 34A). Release of HPMC from LFB lenses containing pre-polymer loaded HPMC released 90% of loaded HPMC in 1 day and reached completion in 2-3 days (FIG. 35). It was clear that the release duration was not sufficient and the LFB formulation had to be altered to increase lens optical clarity (FIG. 36A) by minimizing the swelling of the lens.

The observed swelling was induced by HPMC water absorption. In the LFB lens formulation, the macromer serves as both the oxygen permeable phase and the crosslinking monomer, but due to its large size, the average molecular weight between crosslinks (Mc) is large. The hydrodynamic volume of HPMC in the lens increased as the lens absorbed water and forced the network to expand increasing the volume of the lens. To counter the network expansion, additional crosslinking monomers were added to the formulation to reduce the mesh size and physically limit the expansion of the swollen network by increasing the crosslink density and lowering the Mc. The crosslinking monomers added to the formulation were ethylene glycol dimethacrylate (EGDMA) and polyethylene glycol 200 dimethacrylate (PEG200DMA) at equal weight ratios. These monomers are much shorter than the macromer and restrict the mesh size much more than the macromer. Several other crosslinking monomers were investigated in conjunction with this work but resulted in poor optical clarity and poor modulus values. LFB lenses prepared with EGDMA and PEG200DMA in the formulation are referred to lenses with additional cross linker.

Increasing the crosslinking monomer/HPMC weight ratio (xLer/T) lowers lens swelling (FIG. 34B) and increases the optical clarity of the swollen lenses (FIG. 36B). The optical clarity of lenses synthesized with an HPMC content of 3,000 μg HPMC/lens (120 KDa) and an xLer/T of 1.5 was measured to be 75% transmittance compared to 20% for the lenses prepared without additional cross linkers. Although optical clarity was much improved in swollen lenses containing additional cross linkers, transmittance values were less than an acceptable 90% target and significant haze was observed.

III. Dynamic Drug Release Profile

Release time was also affected by adding additional crosslinking molecules to the formulation. The presence of cross linkers in the lens led to significant improvements in clarity with reduced swelling, but the release times were not extended much and they may not be sufficient to properly control HPMC release for 30 days. By increasing the crosslinking monomer to template ratio (xLer/T) shown in FIG. 37, HPMC release is delayed by the decreased mesh size. As the percentage of cross linker is increased, there is a longer initial delay in release. However, there is not much control over the release rate after the initial delay in release. This is indicated by the equivalent rates of release once approximately 5-10% fractional release is achieved.

To further tailor the HPMC release rate, the molecularly imprinting technique was used within the hydrophilic phase. For HPMC, we used biomimetic analysis of sugar recognition to aid in the choice of functional monomers. Analyzing the amino acid residues of a glucose binding protein yielded valuable information regarding necessary non-covalent interactions. Functional monomers containing carboxylic acids were selected to match corresponding glucose binding protein residue chemistry, such as aspartate, glutamate. DMA was identified as a low affinity functional monomer as the electron delocalization and steric hindrance from the methyl groups resulted in weak and temporary hydrogen bonds. Altering the DMA content in the lens did not significantly alter the release rate, and DMA was not used in the M/T calculation. However, turbidity tests of the formulation and transmittance measurements of lenses with increasing concentration of AA and DMA indicated interaction between the functional monomers and HPMC, resulting in higher transmittance values or lower turbidity. The best explanation is that hydrogen bonding between HPMC and the functional monomers increases the solubility of HPMC in the lens and lens formulations. FIG. 36B highlights the effect of AA on lens clarity.

FIG. 38 demonstrates the varied release rates that can be designed by altering the functional monomer to template (M/T) ratio. Increasing the M/T ratio decreases the release rate and extended the time to complete release from ˜2 days (85% in less than 1 day, control, non-imprinted, no additional cross linker) to 50 days (imprinted lens with M/T ratio of 3.4 containing additional cross linker). It is important to note that the delayed release was not due to the additional crosslinking within the lens. A non-imprinted lens with additional cross linker released all HPMC in less than 10 days. For all lenses with additional cross linker, the crosslinking content was kept approximately constant and the release was significantly decreased as M/T ratio increased. In perfect sink environment, the longest releasing lenses delivered ˜60 μg HPMC/day for 50 days. By extrapolating the rate curve on FIG. 39 indicates that there may be a region where HPMC is sequestered without release due to a molecular Velcro™ effect as demonstrated in our previous work with HA imprinted DAILIES®.

IV. Molecular Imprinting Analysis

One of the traditional ways of demonstrating molecular imprinting is by reloading the template and demonstrating enhanced loading over control polymers. However, given the very high molecular weight of our template molecule (HPMC), this is not a feasible method as all binding will be on the surface of the lens and the template will not enter the bulk structure. In previous work, delayed release of high molecular weight hyaluronic acid was significantly improved through the use of multiple functional monomers, highlighting a synergy between the imprinting monomers providing controlled release. With the significant dependence of release on M/T ratio at low amounts of functional monomer, we hypothesize that molecular imprinting, and not random hydrogen bonding, is the primary reason for the decreased transport observed.

To match commercially available silicone hydrogel lens thicknesses, it was necessary to reduce the swollen lens thickness to 100 μm. In addition, it was desirable to reduce thickness to potentially increase the clarity of the lenses. However, decreasing thickness will reduce the HPMC content and increase the release rate. It is well accepted that release rate will be proportional to the thickness, and that thinner lenses will have an increased rate of release compared to a thicker lenses. It was found that decreasing the thickness limited the loading of HPMC to ˜1,000 μg/lens. Optical clarity was measured on 100 μm thick lenses at various M/T and xLer/T weight ratios (FIG. 40). FIG. 40 demonstrates to maintain greater than 90% transmittance, which is an acceptable commercial optical clarity value, the design variables xLer/T and M/T must stay between 1.0 and 3.5 and 3 and 6, respectively. Therefore, lenses can be produced with therapeutic amounts of HPMC with high optical clarity for commercial use as contact lenses. These lenses also have a water uptake of ˜30%.

V. Additional Dynamic Drug Release Profiles

The next step of this work was to perform a dynamic release study on the thin lenses. Imprinted lenses were formulated to xLer/T of 1.5 and M/T of 3.5 and a HPMC content of 1,050 μg/lens. This is the same lens composition as the lens exhibiting a 52 day HPMC release duration in FIG. 38, except it contained a third of the HPMC (i.e., an equivalent xLer/T and M/T ratio). HPMC release from the thin lens in a perfect sink environment reached completion in 10-12 days. However, the perfect sink model is the fastest possible release, and release in the eye will be slowed due to the reduced sink and physiological conditions. Within an ocular flow rate, the thin lens released HPMC at a linear rate of ˜16 μg HPMC/day for 60 days (FIG. 41). The microfluidic device pioneered by our laboratory models the physiological flow rate of the eye (i.e., 3 μL/min) and was used to more accurately predict the in-vivo lens release profile.

Levels of comfort are not well defined in the literature. Since so little work has been done with continuous release of comfort agents, there are not generally accepted values defining comfort. Until such data is available, estimates must be made from eye drop formulations. A typical over-the-counter HPMC eye drop delivers 0.125 μg HPMC/drop (assuming a 25 μL drop with 0.2%-0.5% re-wetting agent concentration). Drops are applied as needed and 20 drops a day yields a projected cumulative release rate of 2.5 μg HPMC/day. The daily mass delivery rate from the HPMC-imprinted lenses is 6 fold higher than this projected eye drop release rate. Some artificial tears, such as MiniDrops Eye Therapy (Optics Laboratory, Inc.) deliver PVP and PVC at a combined 500 μg comfort agent/drop (20 μg comfort agent/4 assuming a 25 μL drop). With only 2 drops a day, the projected cumulative release rate is 1,000 μg comfort agent/day. Thus, a large range of comfort agent values exist in commercially available products, and our lenses can be designed with varying release rates to alleviate contact lens induced dry eye (CLIDE) symptoms. Also, non-corrective, cosmetic HPMC-imprinted lenses could be used to treat other dry eye conditions. With the imprinted lenses, the HPMC release rate and concentration of comfort molecule in the eye would be relatively constant rather than pulsatile with eye drops. In addition, lenses would need to be replaced on a monthly or weekly basis whereas eye drops might need to be applied each hour or as needed.

The lenses presented in this work can be optimized for various release rates. For example, as shown in FIG. 38, lowering the M/T ratio will increase the release rate, and increasing it will decrease the release rate. The HPMC molecular weight also will have an effect on the release rate with higher molecular weight releasing slower. In addition, other functional monomers interact differently with HPMC and different rates of release can be achieved by altering the imprinting design.

Modulus values were also determined which depended strongly on the xLer/T ratio. Modulus values for the lenses were as follows: (i) LFB lenses: 1.48±0.25 MPa, (ii) LFB lenses prepared with additional cross linkers and no HPMC (equivalent amount of cross linker):12±9 MPa and (iii) LFB lenses with additional cross linkers (xLer/T 1.5) and 1,000 μg HPMC/lens: 8.9±4.3 MPa. As expected, the modulus of the cross-linked lenses increased with additional crosslinking content and decreased when HPMC was added. The modulus can be reduced by small decreases in XLer/T, and small variations in M/T ratio and XLer/T will lead to large variations in release rate, while still producing an optically clear lens. A decrease in M/T and XLer/T will increase the release rate, and an increase in both of these parameters will decrease the release rate.

CONCLUSION

Molecular imprinting has been shown to be a versatile and valuable method of controlling comfort molecule release from soft, extended wear, silicone hydrogel contact lenses. By adding functional monomers and crosslinking monomers, considerable control has been demonstrated over HPMC release rate, optical clarity, swelling, and mechanical properties. Producing a commercial lens prepared with HPMC added to the formulation resulted in a lens with low optical clarity and a fast release of HPMC of insignificant duration compared to wear time. There was no control over release, and poor optical clarity made the lens unsuitable for commercial use. By applying the principles of molecular imprinting, release rates could be significantly varied by the imprinting effect. Release rates can be tailored to deliver 1,000 μg HPMC over a period of up to 60 days in a constant manner. This highlights the tailorable, extended release of a macromolecular re-wetting agent from a silicone hydrogel, extended wear lens. This lens combination device is ideal for combating contact lens induced dry eye and is expected to significantly influence the research and development of future contact lenses as the imprinting technique can be applied as a platform technology to other comfort molecules.

The present invention has been described in terms of specific embodiments incorporating details to facilitate the understanding of the principles of construction and operation of the invention. Such reference herein to specific embodiments and details thereof is not intended to limit the scope of the claims appended hereto. It will be apparent to those skilled in the art that modifications can be made in the embodiment chosen for illustration without departing from the spirit and scope of the invention. Specifically, it will be apparent to one of ordinary skill in the art that the device of the present invention could be implemented in several different ways and the apparatus disclosed above is only illustrative of the preferred embodiment of the invention and is in no way a limitation. 

What is claimed is:
 1. A comfort agent delivery system comprising: a contact lens, the contact lens comprising a recognitive weakly cross-linked polymeric hydrogel matrix with complexing memory sites that complex a comfort agent and release the comfort agent from the hydrogel matrix over time while in contact with a surface of an eye, wherein the weakly cross-linked polymeric hydrogel is formed by the steps of generating a solution comprising amounts of a bio-template, a functionalized monomer or macromer, and a cross-linking monomer, complexing the functionalized monomer or macromer and the bio-template through non-covalent interactions, initiating copolymerization of the functionalized monomer or macromer and the cross-linking monomer, and loading the comfort agent into the memory site, wherein the release rate of the contact lenses is adjusted by either altering the ratio of the functionalized monomer(s) or macromere(s) to the biotemplate, or altering the molecular weight of the comfort agent and/or a drug, wherein the weakly cross-linked polymeric hydrogel provides controlled release of the comfort agent and optionally the drug for at least 50 days.
 2. The composition of claim 1, wherein the contact lenses having the complexing memory sites are produced by the process comprising: a) identifying receptor sites at a target biological tissue that are associated with a biological mechanism of a comfort agent and/or drug at the target biological tissue; b) synthesizing or selecting the functionalized monomer with functional groups that mimic the receptor sites of the target biological tissue to form the matrix for binding the comfort agent and/or drug on the functionalized monomer within a weakly cross-linked polymeric hydrogel; c) forming the polymeric hydrogel comprising: forming a solution comprising amounts of bio-template, the functionalized monomer and a cross-linking monomer; complexing the functionalized monomer and the bio-template in the solution through non-covalent interactions, and initiating copolymerization of the functionalized monomer and the cross-linking monomer; d) forming the weakly cross-linked polymeric hydrogel into contact lens; e) loading the contact lens with the comfort agent and/or drug, wherein the drug complexes with the functionalized monomer as a result of the monomer having the functional groups mimicking the receptor sites at the biological tissue wherein the drug is identified as having a biological mechanism of action.
 3. The composition of claim 1, wherein the ratio of the functionalized monomer to the biotemplate is increased to provide an increased controlled release rate.
 4. The composition of claim 1, wherein the ratio of the functionalized monomer(s) to the biotemplate is at least
 2. 5. The composition of claim 1, wherein at least about 500 μg of the comfort agent and/or drug is delivered over a period of up to 60 days in a constant manner.
 6. The composition of claim 5, wherein at least about 1000 μg of the comfort agent and/or drug is delivered over a period of up to 60 days in a constant manner.
 7. The composition of claim 1, wherein the comfort agent is a lubricating agent useful for treating keratoconjunctivitis sicca (dry eye) and making contact lenses more comfortable.
 8. The composition of claim 1, wherein the hydrogel matrix comprises silicone-based or silicone-containing macromer or polymer chains.
 9. The composition of claim 1, wherein the complexing sites comprise amino acid functional groups.
 10. The composition of claim 1, wherein the contact lens further includes a drug.
 11. The composition of claim 10, wherein the drug is selected from the group consisting of an antibiotic, an anti-inflammatory, an antihistamine, an antiviral agent, a cancer drug, an anesthetic, a cycloplegic, a mydriatics, a lubricant agent, a hydrophilic agent, a decongestant, a vasoconstrictor, vasodilater, an immuno-suppressant, an immuno-modulating agent, an anti-glaucoma agent, an anti-infective, hyperosmolar agent, vitamins, growth factors, growth factor antagonists, sympathomimetics, an adrenergic agonist, an anti-cataract agent, an anti-hypertensive agent, an anti-macular degeneration agent, an ocular permeation enhancing agent, an anti-retinal disease agent, an anti-retinitis pigmentosa agent, an anti-diabetic retinopathy agent, an ocular diagnostic agent, and combinations thereof.
 12. The composition of claim 1, wherein the weakly cross-linked hydrogel matrix comprises: carbon-based polymers or organic-based macromers selected from the group consisting of Polyethylene glycol (200) dimethacrylate (PEG200DMA), ethylene glycol dimethacrylate (EGDMA), tetraethyleneglycol dimethacrylate (TEGDMA), N,N′-Methylene-bis-acrylamide, polyethylene glycol (600) dimethacrylate (PEG600DMA) and combinations thereof; or silicone-based monomers or macromers selected from the group consisting of polydimethyl siloxane-based monomer, tris(trimethylsiloxy)silyl propyl methacrylate (TRIS) and combinations thereof; or hydrophilic TRIS derivatives selected from the group consisting of tris(trimethylsiloxy)silyl propyl vinyl carbamate (TPVC), tris(trimethylsiloxy)silyl propyl glycerol methacrylate (SIGMA), tris(trimethylsiloxy)silyl propyl methacryloxyethylcarbamate (TSMC), polydimethylsiloxane (PDMS) and combinations thereof; or monomers or macromers with pendent silicone groups selected from the group consisting of methacrylate end-capped fluoro-grafted PDMS cross linker, a methacrylate end-capped urethane-siloxane copolymer cross linker, a styrene-capped siloxane polymer containing polyethylene oxide and polypropylene oxide blocks, siloxane containing hydrophilic grafts or amino acid residue grafts, siloxanes containing hydrophilic blocks or containing amino acid residue grafts, and combinations thereof. 